Low-field magnetic resonance imaging methods and apparatus

ABSTRACT

According to some aspects, a low-field magnetic resonance imaging system is provided. The low-field magnetic resonance imaging system comprises a magnetics system having a plurality of magnetics components configured to produce magnetic fields for performing magnetic resonance imaging, the magnetics system comprising, a B 0  magnet configured to produce a B 0  field for the magnetic resonance imaging system at a low-field strength of less than 0.2 Tesla (T), a plurality of gradient coils configured to, when operated, generate magnetic fields to provide spatial encoding of magnetic resonance signals, and at least one radio frequency coil configured to, when operated, transmit radio frequency signals to a field of view of the magnetic resonance imaging system and to respond to magnetic resonance signals emitted from the field of view, a power system comprising one or more power components configured to provide power to the magnetics system to operate the magnetic resonance imaging system to perform image acquisition, and a power connection configured to connect to a single-phase outlet to receive mains electricity and deliver the mains electricity to the power system to provide power needed to operate the magnetic resonance imaging system. According to some aspects, the power system operates the low-field magnetic resonance imaging system using an average of less than 1.6 kilowatts during image acquisition.

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims priority under 35 U.S.C. § 120 and is acontinuation-in-part (CIP) of U.S. application Ser. No. 15/640,369,filed Jun. 30, 2017 and titled “Low-Field Magnetic Resonance ImagingMethods and Apparatus,” which claims priority under 35 U.S.C. § 119 toU.S. Provisional Application Ser. No. 62/425,465, filed Nov. 22, 2016,and titled LOW-FIELD MAGNETIC RESONANCE IMAGING METHODS AND APPARATUS,and this application claims priority under 35 U.S.C. § 119 to U.S.Provisional Application Ser. No. 62/425,465, filed Nov. 22, 2016, andtitled LOW-FIELD MAGNETIC RESONANCE IMAGING METHODS AND APPARATUS, eachapplication of which is herein incorporated by reference in itsentirety.

BACKGROUND

Magnetic resonance imaging (MRI) provides an important imaging modalityfor numerous applications and is widely utilized in clinical andresearch settings to produce images of the inside of the human body. Asa generality, MRI is based on detecting magnetic resonance (MR) signals,which are electromagnetic waves emitted by atoms in response to statechanges resulting from applied electromagnetic fields. For example,nuclear magnetic resonance (NMR) techniques involve detecting MR signalsemitted from the nuclei of excited atoms upon the re-alignment orrelaxation of the nuclear spin of atoms in an object being imaged (e.g.,atoms in the tissue of the human body). Detected MR signals may beprocessed to produce images, which in the context of medicalapplications, allows for the investigation of internal structures and/orbiological processes within the body for diagnostic, therapeutic and/orresearch purposes.

MRI provides an attractive imaging modality for biological imaging dueto the ability to produce non-invasive images having relatively highresolution and contrast without the safety concerns of other modalities(e.g., without needing to expose the subject to ionizing radiation,e.g., x-rays, or introducing radioactive material to the body).Additionally, MRI is particularly well suited to provide soft tissuecontrast, which can be exploited to image subject matter that otherimaging modalities are incapable of satisfactorily imaging. Moreover, MRtechniques are capable of capturing information about structures and/orbiological processes that other modalities are incapable of acquiring.However, there are a number of drawbacks to MRI that, for a givenimaging application, may involve the relatively high cost of theequipment, limited availability and/or difficulty in gaining access toclinical MRI scanners and/or the length of the image acquisitionprocess.

The trend in clinical MRI has been to increase the field strength of MRIscanners to improve one or more of scan time, image resolution, andimage contrast, which, in turn, continues to drive up costs. The vastmajority of installed MRI scanners operate at 1.5 or 3 tesla (T), whichrefers to the field strength of the main magnetic field B₀. A rough costestimate for a clinical MRI scanner is approximately one million dollarsper tesla, which does not factor in the substantial operation, service,and maintenance costs involved in operating such MRI scanners.

Additionally, conventional high-field MRI systems typically requirelarge superconducting magnets and associated electronics to generate astrong uniform static magnetic field (B₀) in which an object (e.g., apatient) is imaged. The size of such systems is considerable with atypical MRI installment including multiple rooms for the magnet,electronics, thermal management system, and control console areas. Thesize and expense of MRI systems generally limits their usage tofacilities, such as hospitals and academic research centers, which havesufficient space and resources to purchase and maintain them. The highcost and substantial space requirements of high-field MRI systemsresults in limited availability of MRI scanners. As such, there arefrequently clinical situations in which an MRI scan would be beneficial,but due to one or more of the limitations discussed above, is notpractical or is impossible, as discussed in further detail below.

SUMMARY

Some embodiments include a low-field magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a B0 magnet configured toproduce a B0 field for the magnetic resonance imaging system at alow-field strength of less than 0.2 Tesla (T), a plurality of gradientcoils configured to, when operated, generate magnetic fields to providespatial encoding of magnetic resonance signals, and at least one radiofrequency coil configured to, when operated, transmit radio frequencysignals to a field of view of the magnetic resonance imaging system andto respond to magnetic resonance signals emitted from the field of view.The low-field magnetic resonance system further comprises a power systemcomprising one or more power components configured to provide power tothe magnetics system to operate the magnetic resonance imaging system toperform image acquisition, and a power connection configured to connectto a single-phase outlet to receive mains electricity and deliver themains electricity to the power system to provide power needed to operatethe magnetic resonance imaging system.

Some embodiments include a low-field magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a B0 magnet configured toproduce a B0 field for the magnetic resonance imaging system at alow-field strength of less than 0.2 Tesla (T), a plurality of gradientcoils configured to, when operated, generate magnetic fields to providespatial encoding of emitted magnetic resonance signals, and at least oneradio frequency coil configured to, when operated, transmit radiofrequency signals to a field of view of the magnetic resonance imagingsystem and to respond to magnetic resonance signals emitted from thefield of view, and a power system comprising one or more powercomponents configured to provide power to the magnetics system tooperate the magnetic resonance imaging system to perform imageacquisition, wherein the power system operates the low-field magneticresonance imaging system using an average of less than 5 kilowattsduring image acquisition.

Some embodiments include a low-field magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a B₀ magnet configured toproduce a B₀ field for the magnetic resonance imaging system, aplurality of gradient coils configured to, when operated, generatemagnetic fields to provide spatial encoding of emitted magneticresonance signals, and at least one radio frequency coil configured to,when operated, transmit radio frequency signals to the field of view ofthe magnetic resonance imaging system and to respond to magneticresonance signals emitted from the field of view. The low-field magneticresonance imaging system further comprises a power system comprising oneor more power components configured to provide power to the magneticssystem to operate the magnetic resonance imaging system to perform imageacquisition, wherein the power system operates the low-field magneticresonance imaging system using an average of less than 1.6 kilowattsduring image acquisition.

Some embodiments include a portable magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a permanent B0 magnetconfigured to produce a B0 field for the magnetic resonance imagingsystem and a plurality of gradient coils configured to, when operated,generate magnetic fields to provide spatial encoding of emitted magneticresonance signals. The portable magnetic resonance imaging systemfurther comprises a power system comprising one or more power componentsconfigured to provide power to the magnetics system to operate themagnetic resonance imaging system to perform image acquisition, and abase that supports the magnetics system and houses the power system, thebase comprising at least one conveyance mechanism allowing the portablemagnetic resonance imaging system to be transported to differentlocations.

Some embodiments include a portable magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a permanent B₀ magnetconfigured to produce a B₀ field for the magnetic resonance imagingsystem, a plurality of gradient coils configured to, when operated,generate magnetic fields to provide spatial encoding of emitted magneticresonance signals, and at least one radio frequency transmit coil. Theportable magnetic resonance imaging system further comprises powersystem comprising one or more power components configured to providepower to the magnetics system to operate the magnetic resonance imagingsystem to perform image acquisition, and a base that supports themagnetics system and houses the power system, the base having a maximumhorizontal dimension of less than or equal to approximately 50 inches.

Some embodiments include a portable magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a permanent B₀ magnetconfigured to produce a B₀ field for the magnetic resonance imagingsystem, a plurality of gradient coils configured to, when operated,generate magnetic fields to provide spatial encoding of emitted magneticresonance signals, and at least one radio frequency transmit coil. Theportable magnetic resonance imaging system further comprises powersystem comprising one or more power components configured to providepower to the magnetics system to operate the magnetic resonance imagingsystem to perform image acquisition, and a base that supports themagnetics system and houses the power system, wherein the portablemagnetic resonance imaging system weighs less than 1,500 pounds.

Some embodiments include a low-field magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a permanent B₀ magnetconfigured to produce a B₀ field having a field strength of less than orequal to approximately 0.1 T, and a plurality of gradient coilsconfigured to, when operated, generate magnetic fields to providespatial encoding of magnetic resonance signals; and at least one radiofrequency coil configured to, when operated, transmit radio frequencysignals to a field of view of the magnetic resonance imaging system andto respond to magnetic resonance signals emitted from the field of view.The low-field magnetic resonance imaging system further comprises atleast one controller configured to operate the magnetics system inaccordance with a predetermined pulse sequence to acquire at least oneimage.

Some embodiments include a low-field magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a permanent B₀ magnetconfigured to produce a B₀ field having a field strength of less than orequal to approximately 0.1 T, and a plurality of gradient coilsconfigured to, when operated, generate magnetic fields to providespatial encoding of magnetic resonance signals; and at least one radiofrequency coil configured to, when operated, transmit radio frequencysignals to a field of view of the magnetic resonance imaging system andto respond to magnetic resonance signals emitted from the field of view,wherein the low-field magnetic resonance imaging system has a 5-Gaussline that has a maximum dimension of less than or equal to five feet.

Some embodiment include a magnetic resonance imaging system comprising aB₀ magnet configured to produce a B₀ field for the magnetic resonanceimaging system, and a positioning member coupled to the B₀ magnet andconfigured to allow the B₀ magnet to be manually rotated to a pluralityof positions, each of the plurality of positions placing the B₀ magnetat a different angle.

Some embodiments include a portable magnetic resonance imaging systemcomprising a magnetics system having a plurality of magnetics componentsconfigured to produce magnetic fields for performing magnetic resonanceimaging, the magnetics system comprising a B0 magnet configured toproduce a B₀ field for the magnetic resonance imaging system, and aplurality of gradient coils configured to, when operated, generatemagnetic fields to provide spatial encoding of emitted magneticresonance signals. The portable magnetic resonance imaging systemfurther comprises a power system comprising one or more power componentsconfigured to provide power to the magnetics system to operate themagnetic resonance imaging system to perform image acquisition, a basethat supports the magnetics system and houses the power system, the basecomprising at least one conveyance mechanism allowing the portablemagnetic resonance imaging system to be transported desired locations,and a positioning member coupled to the B0 magnet and configured toallow the B0 magnet to be rotated to a desired angle.

Some embodiments include a portable magnetic resonance imaging systemcomprising a B0 magnet configured to produce a B0 field for an imagingregion of the magnetic resonance imaging system, a housing for the B0magnet, and at least one electromagnetic shield adjustably coupled tothe housing to provide electromagnetic shielding for the imaging regionin an amount that is configurable by adjusting the at least oneelectromagnetic shield about the imaging region.

Some embodiments include a portable magnetic resonance imaging systemcomprising a B₀ magnet configured to produce a B₀ magnetic field for animaging region of the magnetic resonance imaging system, a noisereduction system configured to detect and suppress at least someelectromagnetic noise in an operating environment of the portablemagnetic resonance imaging system, and electromagnetic shieldingprovided to attenuate at least some of the electromagnetic noise in theoperating environment of the portable magnetic resonance imaging system,the electromagnetic shielding arranged to shield a fraction of theimaging region of the portable magnetic resonance imaging system.

Some embodiments include a portable magnetic resonance imaging systemcomprising a B₀ magnet configured to produce a B₀ field for an imagingregion of the magnetic resonance imaging system, a noise reductionsystem configured to detect and suppress at least some electromagneticnoise in an operating environment of the portable magnetic resonanceimaging system, and electromagnetic shielding for at least a portion ofthe portable magnetic resonance imaging system, the electromagneticshielding providing substantially no shielding of the imaging region ofthe portable magnetic resonance imaging system.

Some embodiments include portable magnetic resonance imaging systemcomprising a B₀ magnet configured to produce a B₀ field for an imagingregion of the magnetic resonance imaging system, a housing for the B₀magnet, and at least one electromagnetic shield structure adjustablycoupled to the housing to provide electromagnetic shielding for theimaging region in an amount that can be varied by adjusting the at leastone electromagnetic shield structure about the imaging region.

BRIEF DESCRIPTION OF THE DRAWINGS

Various aspects and embodiments of the disclosed technology will bedescribed with reference to the following figures. It should beappreciated that the figures are not necessarily drawn to scale.

FIG. 1 illustrates exemplary components of a magnetic resonance imagingsystem;

FIGS. 2A and 2B illustrate a B₀ magnet comprising a plurality ofelectromagnets, in accordance with some embodiments;

FIG. 3A illustrates a B₀ magnet comprising a plurality of permanentmagnets, in accordance with some embodiments;

FIG. 3B illustrates a top view of an exemplary configuration ofpermanent magnet rings forming, in part, the B₀ magnet illustrated inFIG. 3A;

FIGS. 4A and 4B illustrate an exemplary ring of permanent magnets for aB₀ magnet, in accordance with some embodiments;

FIGS. 5A-C illustrate exemplary dimensions for permanent magnet blocksfor the permanent magnet ring illustrated in FIGS. 4A and 4B, inaccordance with some embodiments;

FIGS. 6A-C illustrate exemplary dimensions for permanent magnet blocksfor the permanent magnet ring illustrated in FIGS. 4A and 4B, inaccordance with some embodiments;

FIGS. 7A-7F illustrate respective portions of an exemplary ring ofpermanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 8A-C illustrate exemplary dimensions for permanent magnet blocksfor an inner sub-ring of the permanent magnet ring illustrated in FIGS.7A-F, in accordance with some embodiments;

FIGS. 9A-C illustrate exemplary dimensions for permanent magnet blocksfor a middle sub-ring of the permanent magnet ring illustrated in FIGS.7A-F, in accordance with some embodiments;

FIGS. 10A-C illustrate exemplary dimensions for permanent magnet blocksfor a outer sub-ring of the permanent magnet ring illustrated in FIGS.7A-F, in accordance with some embodiments;

FIGS. 11A-F illustrate portions of an exemplary ring of permanentmagnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 12A-C illustrate exemplary dimensions for permanent magnet blocksfor an inner sub-ring of the permanent magnet ring illustrated in FIGS.11A-F, in accordance with some embodiments;

FIGS. 13A-C illustrate exemplary dimensions for permanent magnet blocksfor a middle sub-ring of the permanent magnet ring illustrated in FIGS.11A-F, in accordance with some embodiments;

FIGS. 14A-C illustrate exemplary dimensions for permanent magnet blocksfor an outer sub-ring of the permanent magnet ring illustrated in FIGS.11A-F, in accordance with some embodiments;

FIGS. 15A-C illustrate views of an exemplary permanent magnet disk, inaccordance with some embodiments;

FIG. 16 illustrates a B0 magnet comprising a plurality of permanentmagnets, in accordance with some embodiments;

FIG. 17 illustrates a top view of an exemplary configuration ofpermanent magnet rings forming, in part, the B₀ magnet illustrated inFIG. 16;

FIGS. 18A and 18B illustrate an exemplary ring of permanent magnetsegments for a B₀ magnet, in accordance with some embodiments;

FIGS. 18C and 18D illustrate different views of permanent magnetsegments that can be used to form the permanent magnet ring illustratedin FIG. 18E, in accordance with some embodiments;

FIG. 18E illustrates a permanent magnet ring for a B₀ magnet, inaccordance with some embodiments;

FIGS. 18F and 18G illustrate different views of permanent magnetsegments that can be used to form the permanent magnet ring illustratedin FIG. 18H, in accordance with some embodiments;

FIG. 18H illustrates a permanent magnet ring for a B₀ magnet, inaccordance with some embodiments;

FIGS. 19A and 19B illustrate a portable low-field MRI system, inaccordance with some embodiments.

FIG. 20 shows drive circuitry for driving a current through a coil toproduce a magnetic field, in accordance with some embodiments of thetechnology described herein.

FIG. 21A shows an example of a gradient coil current waveform, inaccordance with some embodiments of the technology described herein.

FIG. 21B shows waveforms for the current command, the gradient coilcurrent and the gradient coil voltage before, during and after therising transition of the gradient coil current waveform shown in FIG.21A, in accordance with some embodiments of the technology describedherein.

FIG. 22A shows an example of a power component having a current feedbackloop and a voltage feedback loop, in accordance with some embodiments ofthe technology described herein.

FIG. 22B shows an example of a voltage amplifier, in accordance withsome embodiments of the technology described herein.

FIGS. 23A and 23B show examples of an output stage that can be poweredby different supply terminals depending on the output voltage, inaccordance with some embodiments of the technology described herein.

FIG. 24 shows an example of an output stage having a plurality of drivecircuits to drive a plurality of transistor circuits connected to highvoltage and low voltage supply terminals, in accordance with someembodiments of the technology described herein.

FIG. 25 shows drive circuits including a bias circuit and a timercircuit, in accordance with some embodiments of the technology describedherein.

FIG. 26 shows an example implementation of the drive circuits of FIG.25, in accordance with some embodiments of the technology describedherein.

FIG. 27 shows another example of a technique for implementing a timingcircuit, according to some embodiments.

FIG. 28 shows an example of timing circuits realized by an RC circuitand a transistor, according to some embodiments.

FIG. 29 shows an example of an output stage including a single-endedlinear amplifier, according to some embodiments.

FIG. 30 shows an example of a power component may include a switchingpower converter, according to some embodiments.

FIG. 31 shows an embodiment of an output stage that may be powered by avariable voltage positive supply terminal and a variable voltagenegative supply terminal, according to some embodiments.

FIG. 32A shows an embodiment similar to that of FIG. 23A with variablelow voltage supply terminals.

FIG. 32B shows an embodiment in which the high voltage supply terminalsare the same as the power supply terminals that supply power to thepower converters.

FIGS. 33A-D show a gradient coil current waveform, gradient coil voltagewaveform, and power supply terminal voltage waveforms, according to someembodiments.

FIG. 34A shows an embodiment similar to that of FIG. 30 with a variablelow voltage supply terminal.

FIG. 34B shows an embodiment in which the high voltage supply terminalis the same as the power supply terminal that supplies power to thepower converter.

FIG. 35 illustrates a radio frequency power amplifier (RFPA), inaccordance with some embodiments;

FIG. 36 illustrates a housing for electronic components of a portableMRI system, in accordance with some embodiments;

FIG. 37A illustrates a circular housing for electronic components of aportable MRI system, in accordance with some embodiments;

FIGS. 37B and 37C illustrate views of a base comprising a housing forelectronics components of a portable MRI system, in accordance with someembodiments;

FIG. 37D illustrates a portable MRI system, in accordance with someembodiments;

FIG. 38A illustrates permanent magnet shims for a B₀ magnet of aportable MRI system, in accordance with some embodiments;

FIGS. 38B and 38C illustrate vibration mounts for gradient coils of aportable MRI system, in accordance with some embodiments;

FIG. 38D illustrates a laminate panel comprising gradient coils fastenedto the vibration mounts illustrated in FIGS. 38B and 38C;

FIG. 38E illustrates exemplary shims for a B₀ magnet of a portable MRIsystem, in accordance with some embodiments;

FIG. 38F illustrates a portable MRI system, in accordance with someembodiments;

FIGS. 39A and 39B illustrate views of a portable MRI system, inaccordance with some embodiments;

FIG. 39C illustrates another example of a portable MRI system, inaccordance with some embodiments;

FIG. 40A illustrates a portable MRI system performing a scan of thehead, in accordance with some embodiments;

FIG. 40B illustrates a portable MRI system performing a scan of theknee, in accordance with some embodiments;

FIG. 41A-D illustrate exemplary respective examples of a noise reductionsystem, in accordance with some embodiments;

FIG. 42 is a flowchart of an illustrative process for performing noisereduction, in accordance with some embodiments;

FIG. 43A-B illustrate respective examples of decoupling circuitsconfigured to reduce inductive coupling between radio frequency coils ina multi-coil transmit/receive system, in accordance with someembodiments;

FIG. 43C illustrates a decoupling circuit using Gallium Nitride (GaN)field effect transistors (FETs) to couple and decouple a receive coil,in accordance with some embodiments;

FIG. 43D illustrates an active decoupling circuit, in accordance withsome embodiments;

FIGS. 44A-C illustrate a portable MRI system having different amounts ofdevice-level shielding about the imaging region, in accordance with someembodiments;

FIG. 44D illustrates a portable MRI system utilizing a conductive stripto provide electromagnetic shielding for the imaging region, inaccordance with some embodiments;

FIGS. 45A-45D illustrate different views of a positioning mechanism, inaccordance with some embodiments;

FIGS. 46A and 46B illustrate exemplary components of a positioningmechanism, in accordance with some embodiments; and

FIGS. 47-50 illustrate images obtained using the low-field MRI systemsdescribed herein, in accordance with some embodiments.

DETAILED DESCRIPTION

The MRI scanner market is overwhelmingly dominated by high-fieldsystems, and particularly for medical or clinical MRI applications. Asdiscussed above, the general trend in medical imaging has been toproduce MRI scanners with increasingly greater field strengths, with thevast majority of clinical MRI scanners operating at 1.5 T or 3 T, withhigher field strengths of 7 T and 9 T used in research settings. As usedherein, “high-field” refers generally to MRI systems presently in use ina clinical setting and, more particularly, to MRI systems operating witha main magnetic field (i.e., a B₀ field) at or above 1.5 T, thoughclinical systems operating between 0.5 T and 1.5 T are often alsocharacterized as “high-field.” Field strengths between approximately 0.2T and 0.5 T have been characterized as “mid-field” and, as fieldstrengths in the high-field regime have continued to increase, fieldstrengths in the range between 0.5 T and 1 T have also beencharacterized as mid-field. By contrast, “low-field” refers generally toMRI systems operating with a B₀ field of less than or equal toapproximately 0.2 T, though systems having a B₀ field of between 0.2 Tand approximately 0.3 T have sometimes been characterized as low-fieldas a consequence of increased field strengths at the high end of thehigh-field regime. Within the low-field regime, low-field MRI systemsoperating with a B₀ field of less than 0.1 T are referred to herein as“very low-field” and low-field MRI systems operating with a B₀ field ofless than 10 mT are referred to herein as “ultra-low field.”

As discussed above, conventional MRI systems require specializedfacilities. An electromagnetically shielded room is required for the MRIsystem to operate and the floor of the room must be structurallyreinforced. Additional rooms must be provided for the high-powerelectronics and the scan technician's control area. Secure access to thesite must also be provided. In addition, a dedicated three-phaseelectrical connection must be installed to provide the power for theelectronics that, in turn, are cooled by a chilled water supply.Additional HVAC capacity typically must also be provided. These siterequirements are not only costly, but significantly limit the locationswhere MRI systems can be deployed. Conventional clinical MRI scannersalso require substantial expertise to both operate and maintain. Thesehighly trained technicians and service engineers add large on-goingoperational costs to operating an MRI system. Conventional MRI, as aresult, is frequently cost prohibitive and is severely limited inaccessibility, preventing MRI from being a widely available diagnostictool capable of delivering a wide range of clinical imaging solutionswherever and whenever needed. Typically, patient must visit one of alimited number of facilities at a time and place scheduled in advance,preventing MRI from being used in numerous medical applications forwhich it is uniquely efficacious in assisting with diagnosis, surgery,patient monitoring and the like.

As discussed above, high-field MRI systems require specially adaptedfacilities to accommodate the size, weight, power consumption andshielding requirements of these systems. For example, a 1.5 T MRI systemtypically weighs between 4-10 tons and a 3 T MRI system typically weighsbetween 8-20 tons. In addition, high-field MRI systems generally requiresignificant amounts of heavy and expensive shielding. Many mid-fieldscanners are even heavier, weighing between 10-20 tons due, in part, tothe use of very large permanent magnets and/or yokes. Commerciallyavailable low-field MRI systems (e.g., operating with a B₀ magneticfield of 0.2 T) are also typically in the range of 10 tons or more duethe large of amounts of ferromagnetic material used to generate the B₀field, with additional tonnage in shielding. To accommodate this heavyequipment, rooms (which typically have a minimum size of 30-50 squaremeters) have to be built with reinforced flooring (e.g., concreteflooring), and must be specially shielded to prevent electromagneticradiation from interfering with operation of the MRI system. Thus,available clinical MRI systems are immobile and require the significantexpense of a large, dedicated space within a hospital or facility, andin addition to the considerable costs of preparing the space foroperation, require further additional on-going costs in expertise inoperating and maintaining the system.

In addition, currently available MRI systems typically consume largeamounts of power. For example, common 1.5 T and 3 T MRI systemstypically consume between 20-40 kW of power during operation, whileavailable 0.5 T and 0.2 T MRI systems commonly consume between 5-20 kW,each using dedicated and specialized power sources. Unless otherwisespecified, power consumption is referenced as average power consumedover an interval of interest. For example, the 20-40 kW referred toabove indicates the average power consumed by conventional MRI systemsduring the course of image acquisition, which may include relativelyshort periods of peak power consumption that significantly exceeds theaverage power consumption (e.g., when the gradient coils and/or RF coilsare pulsed over relatively short periods of the pulse sequence).Intervals of peak (or large) power consumption are typically addressedvia power storage elements (e.g., capacitors) of the MRI system itself.Thus, the average power consumption is the more relevant number as itgenerally determines the type of power connection needed to operate thedevice. As discussed above, available clinical MRI systems must havededicated power sources, typically requiring a dedicated three-phaseconnection to the grid to power the components of the MRI system.Additional electronics are then needed to convert the three-phase powerinto single-phase power utilized by the MRI system. The many physicalrequirements of deploying conventional clinical MRI systems creates asignificant problem of availability and severely restricts the clinicalapplications for which MRI can be utilized.

Accordingly, the many requirements of high-field MRI renderinstallations prohibitive in many situations, limiting their deploymentto large institutional hospitals or specialized facilities and generallyrestricting their use to tightly scheduled appointments, requiring thepatient to visit dedicated facilities at times scheduled in advance.Thus, the many restrictions on high field MRI prevent MRI from beingfully utilized as an imaging modality. Despite the drawbacks ofhigh-field MRI mentioned above, the appeal of the significant increasein SNR at higher fields continues to drive the industry to higher andhigher field strengths for use in clinical and medical MRI applications,further increasing the cost and complexity of MRI scanners, and furtherlimiting their availability and preventing their use as ageneral-purpose and/or generally-available imaging solution.

The low SNR of MR signals produced in the low-field regime (particularlyin the very low-field regime) has prevented the development of arelatively low cost, low power and/or portable MRI system. Conventional“low-field” MRI systems operate at the high end of what is typicallycharacterized as the low-field range (e.g., clinically availablelow-field systems have a floor of approximately 0.2 T) to achieve usefulimages. Though somewhat less expensive then high-field MRI systems,conventional low-field MRI systems share many of the same drawbacks. Inparticular, conventional low-field MRI systems are large, fixed andimmobile installments, consume substantial power (requiring dedicatedthree-phase power hook-ups) and require specially shielded rooms andlarge dedicated spaces. The challenges of low-field MRI have preventedthe development of relatively low cost, low power and/or portable MRIsystems that can produce useful images.

The inventors have developed techniques enabling portable, low-field,low power and/or lower-cost MRI systems that can improve the wide-scaledeployability of MRI technology in a variety of environments beyond thecurrent MRI installments at hospitals and research facilities. As aresult, MRI can be deployed in emergency rooms, small clinics, doctor'soffices, in mobile units, in the field, etc. and may be brought to thepatient (e.g., bedside) to perform a wide variety of imaging proceduresand protocols. Some embodiments include very low-field MRI systems(e.g., 0.1 T, 50 mT, 20 mT, etc.) that facilitate portable, low-cost,low-power MRI, significantly increasing the availability of MRI in aclinical setting.

There are numerous challenges to developing a clinical MRI system in thelow-field regime. As used herein, the term clinical MRI system refers toan MRI system that produces clinically useful images, which refers to animages having sufficient resolution and adequate acquisition times to beuseful to a physician or clinician for its intended purpose given aparticular imaging application. As such, the resolutions/acquisitiontimes of clinically useful images will depend on the purpose for whichthe images are being obtained. Among the numerous challenges inobtaining clinically useful images in the low-field regime is therelatively low SNR. Specifically, the relationship between SNR and B₀field strength is approximately B₀ ^(5/4) at field strength above 0.2 Tand approximately B₀ ^(3/2) at field strengths below 0.1 T. As such, theSNR drops substantially with decreases in field strength with even moresignificant drops in SNR experienced at very low field strength. Thissubstantial drop in SNR resulting from reducing the field strength is asignificant factor that has prevented development of clinical MRIsystems in the very low-field regime. In particular, the challenge ofthe low SNR at very low field strengths has prevented the development ofa clinical MRI system operating in the very low-field regime. As aresult, clinical MRI systems that seek to operate at lower fieldstrengths have conventionally achieved field strengths of approximatelythe 0.2 T range and above. These MRI systems are still large, heavy andcostly, generally requiring fixed dedicated spaces (or shielded tents)and dedicated power sources.

The inventors have developed low-field and very low-field MRI systemscapable of producing clinically useful images, allowing for thedevelopment of portable, low cost and easy to use MRI systems notachievable using state of the art technology. According to someembodiments, an MRI system can be transported to the patient to providea wide variety of diagnostic, surgical, monitoring and/or therapeuticprocedures, generally, whenever and wherever needed.

FIG. 1 is a block diagram of typical components of a MRI system 100. Inthe illustrative example of FIG. 1, MRI system 100 comprises computingdevice 104, controller 106, pulse sequences store 108, power managementsystem 110, and magnetics components 120. It should be appreciated thatsystem 100 is illustrative and that a MRI system may have one or moreother components of any suitable type in addition to or instead of thecomponents illustrated in FIG. 1. However, a MRI system will generallyinclude these high level components, though the implementation of thesecomponents for a particular MRI system may differ vastly, as discussedin further detail below.

As illustrated in FIG. 1, magnetics components 120 comprise B₀ magnet122, shim coils 124, RF transmit and receive coils 126, and gradientcoils 128. Magnet 122 may be used to generate the main magnetic fieldB₀. Magnet 122 may be any suitable type or combination of magneticscomponents that can generate a desired main magnetic B₀ field. Asdiscussed above, in the high field regime, the B₀ magnet is typicallyformed using superconducting material generally provided in a solenoidgeometry, requiring cryogenic cooling systems to keep the B₀ magnet in asuperconducting state. Thus, high-field B₀ magnets are expensive,complicated and consume large amounts of power (e.g., cryogenic coolingsystems require significant power to maintain the extremely lowtemperatures needed to keep the B₀ magnet in a superconducting state),require large dedicated spaces, and specialized, dedicated powerconnections (e.g., a dedicated three-phase power connection to the powergrid). Conventional low-field B₀ magnets (e.g., B₀ magnets operating at0.2 T) are also often implemented using superconducting material andtherefore have these same general requirements. Other conventionallow-field B₀ magnets are implemented using permanent magnets, which toproduce the field strengths to which conventional low-field systems arelimited (e.g., between 0.2 T and 0.3 T due to the inability to acquireuseful images at lower field strengths), need to be very large magnetsweighing 5-20 tons. Thus, the B₀ magnet of conventional MRI systemsalone prevents both portability and affordability.

Gradient coils 128 may be arranged to provide gradient fields and, forexample, may be arranged to generate gradients in the B₀ field in threesubstantially orthogonal directions (X, Y, Z). Gradient coils 128 may beconfigured to encode emitted MR signals by systematically varying the B₀field (the B₀ field generated by magnet 122 and/or shim coils 124) toencode the spatial location of received MR signals as a function offrequency or phase. For example, gradient coils 128 may be configured tovary frequency or phase as a linear function of spatial location along aparticular direction, although more complex spatial encoding profilesmay also be provided by using nonlinear gradient coils. For example, afirst gradient coil may be configured to selectively vary the B₀ fieldin a first (X) direction to perform frequency encoding in thatdirection, a second gradient coil may be configured to selectively varythe B₀ field in a second (Y) direction substantially orthogonal to thefirst direction to perform phase encoding, and a third gradient coil maybe configured to selectively vary the B₀ field in a third (Z) directionsubstantially orthogonal to the first and second directions to enableslice selection for volumetric imaging applications. As discussed above,conventional gradient coils also consume significant power, typicallyoperated by large, expensive gradient power sources, as discussed infurther detail below.

MRI is performed by exciting and detecting emitted MR signals usingtransmit and receive coils, respectively (often referred to as radiofrequency (RF) coils). Transmit/receive coils may include separate coilsfor transmitting and receiving, multiple coils for transmitting and/orreceiving, or the same coils for transmitting and receiving. Thus, atransmit/receive component may include one or more coils fortransmitting, one or more coils for receiving and/or one or more coilsfor transmitting and receiving. Transmit/receive coils are also oftenreferred to as Tx/Rx or Tx/Rx coils to generically refer to the variousconfigurations for the transmit and receive magnetics component of anMRI system. These terms are used interchangeably herein. In FIG. 1, RFtransmit and receive coils 126 comprise one or more transmit coils thatmay be used to generate RF pulses to induce an oscillating magneticfield B₁. The transmit coil(s) may be configured to generate anysuitable types of RF pulses.

Power management system 110 includes electronics to provide operatingpower to one or more components of the low-field MRI system 100. Forexample, as discussed in more detail below, power management system 110may include one or more power supplies, gradient power components,transmit coil components, and/or any other suitable power electronicsneeded to provide suitable operating power to energize and operatecomponents of MRI system 100. As illustrated in FIG. 1, power managementsystem 110 comprises power supply 112, power component(s) 114,transmit/receive switch 116, and thermal management components 118(e.g., cryogenic cooling equipment for superconducting magnets). Powersupply 112 includes electronics to provide operating power to magneticcomponents 120 of the MRI system 100. For example, power supply 112 mayinclude electronics to provide operating power to one or more B₀ coils(e.g., B₀ magnet 122) to produce the main magnetic field for thelow-field MRI system. Transmit/receive switch 116 may be used to selectwhether RF transmit coils or RF receive coils are being operated.

Power component(s) 114 may include one or more RF receive (Rx)pre-amplifiers that amplify MR signals detected by one or more RFreceive coils (e.g., coils 126), one or more RF transmit (Tx) powercomponents configured to provide power to one or more RF transmit coils(e.g., coils 126), one or more gradient power components configured toprovide power to one or more gradient coils (e.g., gradient coils 128),and one or more shim power components configured to provide power to oneor more shim coils (e.g., shim coils 124).

In conventional MRI systems, the power components are large, expensiveand consume significant power. Typically, the power electronics occupy aroom separate from the MRI scanner itself. The power electronics notonly require substantial space, but are expensive complex devices thatconsume substantial power and require wall mounted racks to besupported. Thus, the power electronics of conventional MRI systems alsoprevent portability and affordable of MRI.

As illustrated in FIG. 1, MRI system 100 includes controller 106 (alsoreferred to as a console) having control electronics to sendinstructions to and receive information from power management system110. Controller 106 may be configured to implement one or more pulsesequences, which are used to determine the instructions sent to powermanagement system 110 to operate the magnetic components 120 in adesired sequence (e.g., parameters for operating the RF transmit andreceive coils 126, parameters for operating gradient coils 128, etc.).As illustrated in FIG. 1, controller 106 also interacts with computingdevice 104 programmed to process received MR data. For example,computing device 104 may process received MR data to generate one ormore MR images using any suitable image reconstruction process(es).Controller 106 may provide information about one or more pulse sequencesto computing device 104 for the processing of data by the computingdevice. For example, controller 106 may provide information about one ormore pulse sequences to computing device 104 and the computing devicemay perform an image reconstruction process based, at least in part, onthe provided information. In conventional MRI systems, computing device104 typically includes one or more high performance work-stationsconfigured to perform computationally expensive processing on MR datarelatively rapidly. Such computing devices are relatively expensiveequipment on their own.

As should be appreciated from the foregoing, currently availableclinical MRI systems (including high-field, mid-field and low-fieldsystems) are large, expensive, fixed installations requiring substantialdedicated and specially designed spaces, as well as dedicated powerconnections. The inventors have developed low-field, including very-lowfield, MRI systems that are lower cost, lower power and/or portable,significantly increasing the availability and applicability of MRI.According to some embodiments, a portable MRI system is provided,allowing an MRI system to be brought to the patient and utilized atlocations where it is needed.

As discussed above, some embodiments include an MRI system that isportable, allowing the MRI device to be moved to locations in which itis needed (e.g., emergency and operating rooms, primary care offices,neonatal intensive care units, specialty departments, emergency andmobile transport vehicles and in the field). There are numerouschallenges that face the development of a portable MRI system, includingsize, weight, power consumption and the ability to operate in relativelyuncontrolled electromagnetic noise environments (e.g., outside aspecially shielded room). As discussed above, currently availableclinical MRI systems range from approximately 4-20 tons. Thus, currentlyavailable clinical MRI systems are not portable because of the sheersize and weight of the imaging device itself, let alone the fact thatcurrently available systems also require substantial dedicated space,including a specially shielded room to house the MRI scanner andadditional rooms to house the power electronics and the techniciancontrol area, respectively. The inventors have developed MRI systems ofsuitable weight and size to allow the MRI system to be transported to adesired location, some examples of which are discussed in further detailbelow.

A further aspect of portability involves the capability of operating theMRI system in a wide variety of locations and environments. As discussedabove, currently available clinical MRI scanners are required to belocated in specially shielded rooms to allow for correct operation ofthe device and is one (among many) of the reasons contributing to thecost, lack of availability and non-portability of currently availableclinical MRI scanners. Thus, to operate outside of a specially shieldedroom and, more particularly, to allow for generally portable, cartableor otherwise transportable MRI, the MRI system must be capable ofoperation in a variety of noise environments. The inventors havedeveloped noise suppression techniques that allow the MRI system to beoperated outside of specially shielded rooms, facilitating bothportable/transportable MRI as well as fixed MRI installments that do notrequire specially shielded rooms. While the noise suppression techniquesallow for operation outside specially shielded rooms, these techniquescan also be used to perform noise suppression in shielded environments,for example, less expensive, loosely or ad-hoc shielding environments,and can be therefore used in conjunction with an area that has beenfitted with limited shielding, as the aspects are not limited in thisrespect.

A further aspect of portability involves the power consumption of theMRI system. As also discussed above, current clinical MRI systemsconsume large amounts of power (e.g., ranging from 20 kW to 40 kWaverage power consumption during operation), thus requiring dedicatedpower connections (e.g., dedicated three-phase power connections to thegrid capable of delivering the required power). The requirement of adedicated power connection is a further obstacle to operating an MRIsystem in a variety of locations other than expensive dedicated roomsspecially fitted with the appropriate power connections. The inventorshave developed low power MRI systems capable of operating using mainselectricity such as a standard wall outlet (e.g., 120V/20 A connectionin the U.S.) or common large appliance outlets (e.g., 220-240V/30 A),allowing the device to be operated anywhere common power outlets areprovided. The ability to “plug into the wall” facilitates bothportable/transportable MRI as well as fixed MRI system installationswithout requiring special, dedicated power such as a three-phase powerconnection.

According to some embodiments, a portable MRI system (e.g., any of theportable MRI systems illustrated in FIGS. 19, 39-40 and 44A-D below) isconfigured to operate using mains electricity (e.g., single-phaseelectricity provided at standard wall outlets) via a power connection3970 (see e.g., FIG. 39B). According to some embodiments, a portable MRIsystem comprises a power connection configured to connect to asingle-phase outlet providing approximately between 110 and 120 voltsand rated at 15, 20 or 30 amperes, and wherein the power system iscapable of providing the power to operate the portable MRI system frompower provided by the single-phase outlet. According to someembodiments, a portable MRI system comprises a power connectionconfigured to connect to a single-phase outlet providing approximatelybetween 220 and 240 volts and rated at 15, 20 or 30 amperes, and whereinthe power system is capable of providing the power to operate themagnetic resonance imaging system from power provided by thesingle-phase outlet. According to some embodiments, a portable MRIsystem is configured using the low power techniques described herein touse an average of less than 3 kilowatts during image acquisition.According to some embodiments, a portable MRI system is configured usingthe low power techniques described herein to use an average of less than2 kilowatts during image acquisition. According to some embodiments, aportable MRI system is configured using the low power techniquesdescribed herein to use an average of less than 1 kilowatt during imageacquisition. For example, a low power MRI system employing a permanentB₀ magnet and low power components described herein may operate at 1kilowatt or less, such as at 750 watts or less.

As discussed above, a significant contributor to the size, cost andpower consumption of conventional MRI systems are the power electronicsfor powering the magnetics components of the MRI system. The powerelectronics for conventional MRI systems often require a separate room,are expensive and consume significant power to operate the correspondingmagnetics components. In particular, the gradient coils and thermalmanagement systems utilized to cool the gradient coils alone generallyrequire dedicated power connections and prohibit operation from standardwall outlets. The inventors have developed low power, low noise gradientpower sources capable of powering the gradient coils of an MRI systemthat can, in accordance with some embodiments, be housed in the sameportable, cartable or otherwise transportable apparatus as the magneticscomponents of the MRI system. According to some embodiments, the powerelectronics for powering the gradient coils of an MRI system consumeless than 50 W when the system is idle and between 100-200 W when theMRI system is operating (i.e., during image acquisition). The inventorshave developed power electronics (e.g., low power, low noise powerelectronics) to operate a portable low field MRI system that all fitwithin the footprint of the portable MRI scanner. According to someembodiments, innovative mechanical design has enabled the development ofan MRI scanner that is maneuverable within the confines of a variety ofclinical environments in which the system is needed.

At the core of developing a low power, low cost and/or portable MRIsystem is the reduction of the field strength of the B₀ magnet, whichcan facilitate a reduction in size, weight, expense and powerconsumption. However, as discussed above, reducing the field strengthhas a corresponding and significant reduction in SNR. This significantreduction in SNR has prevented clinical MRI systems from reducing thefield strength below the current floor of approximately 0.2 T, whichsystems remains large, heavy, expensive fixed installations requiringspecialized and dedicated spaces. While some systems have been developedthat operate between 0.1 T and 0.2 T, these systems are oftenspecialized devices for scanning extremities such as the hand, arm orknee. The inventors have developed MRI systems operating in thelow-field and very-low field capable of acquiring clinically usefulimages. Some embodiments include highly efficient pulse sequences thatfacilitate acquiring clinically useful images at lower field strengthsthan previously achievable. The signal to noise ratio of the MR signalis related to the strength of the main magnetic field B₀, and is one ofthe primary factors driving clinical systems to operate in thehigh-field regime. Pulse sequences developed by the inventors thatfacilitate acquisition of clinically useful images are described in U.S.patent application Ser. No. 14/938,430, filed Nov. 11, 2015 and titled“Pulse Sequences for Low Field Magnetic Resonance,” which is hereinincorporated by reference in its entirety.

Further techniques developed by the inventors to address the low SNR oflow field strength include optimizing the configuration of radiofrequency (RF) transmit and/or receive coils to improve the ability ofthe RF transmit/receive coils to transmit magnetic fields and detectemitted MR signals. The inventors have appreciated that the low transmitfrequencies in the low field regime allow for RF coil designs notpossible at higher fields strengths and have developed RF coils withimproved sensitivity, thereby increasing the SNR of the MRI system.Exemplary RF coil designs and optimization techniques developed by theinventors are described in U.S. patent application Ser. No. 15/152,951,filed May 12, 2016 and titled “Radio Frequency Coil Methods andApparatus,” which is herein incorporated by reference in its entirety.

Another technique for addressing the relatively low SNR characteristicof the low-field regime is to improve the homogeneity of the B₀ field bythe B₀ magnet. In general, a B₀ magnet requires some level of shimmingto produce a B₀ magnetic field with a profile (e.g., a B₀ magnetic fieldat the desired field strength and/or homogeneity) satisfactory for usein MRI. In particular, production factors such as design, manufacturingtolerances, imprecise production processes, environment, etc., give riseto field variation that produces a B₀ field having unsatisfactoryprofile after assembly/manufacture. For example, after production,exemplary B₀ magnets 200, 300 and/or 3200 described above may produce aB₀ field with an unsatisfactory profile (e.g., inhomogeneity in the B₀field unsuitable for imaging) that needs to be improved or otherwisecorrected, typically by shimming, to produce clinically useful images.Shimming refers to any of various techniques for adjusting, correctingand/or improving a magnetic field, often the B₀ magnetic field of amagnetic resonance imaging device. Similarly, a shim refers to something(e.g., an object, component, device, system or combination thereof) thatperforms shimming (e.g., by producing a magnetic field).

Conventional techniques for shimming are relatively time and/or costintensive, often requiring significant manual effort by an expert inorder to adjust the B₀ magnetic field so that is it suitable for itsintended purpose, which incurs significant post-production time andexpense. For example, conventional shimming techniques typically involvean iterative process by which the B₀ magnetic field is measured, thenecessary corrections are determined and deployed, and the processrepeated until a satisfactory B₀ magnetic field is produced. Thisiterative process is conventionally performed with substantial manualinvolvement, requiring expertise and significant time (e.g., a day at aminimum, and more typically, longer). Thus, conventional post-productionfield correction of a B₀ magnetic field significantly contributes to theexpense and complexity of conventional MRI systems.

The inventors have developed a number of techniques that, according tosome embodiments, facilitate more efficient and/or cost effectiveshimming for a B₀ magnet for MRI. Some embodiments are suitable for usein low-field MRI, but the techniques described herein are not limitedfor use in the low-field context. For example, the inventors havedeveloped techniques to minimize the manual effort involved incorrecting the B₀ field produced by a B₀ magnet, for example, correctingat least some field inhomogeneity resulting from imperfect manufacturingprocesses. In particular, the inventors have developed automatedtechniques for patterning magnetic material to provide accurate andprecise field correction to the B₀ field produced by a B₀ magnet.Exemplary shimming techniques developed by the inventors are describedin U.S. patent application Ser. No. 15/466,500, filed Mar. 22, 2017 andtitled “Methods and Apparatus for Magnetic Field Shimming,” which isherein incorporated by reference in its entirety.

Another aspect of increasing the availability of MRI is to make MRIaffordable. The development of a portable low-field MRI system by theinventors eliminates many of the costs associated with conventionalclinical MRI systems, including expensive superconducting materials andcryogenic cooling systems, expensive site preparation of large andcomplex dedicated facilities, highly trained personnel to operate andmaintain the system to name a few. In addition, the inventors havedeveloped further cost reduction techniques and designs, including,according to some embodiments, integrated power electronics, designsthat reduce materials, optimize or otherwise minimize the use ofexpensive materials and/or reduce production costs. The inventors havedeveloped automated shimming techniques to allow for correction of fieldinhomogeneity of the B₀ magnet after manufacture, reducing the cost ofboth production and post-production processes.

According to some embodiments, designs developed by the inventors alsoreduce the cost and complexity of operating and maintaining the MRIscanner. For example, conventional clinical MRI systems requiresignificant expertise to both operate and maintain, resulting insignificant on-going costs of these systems. The inventors havedeveloped easy-to-use an MRI systems that allow minimally trained oruntrained personnel to operate and/or maintain the system. According tosome embodiments, automated setup processes allow the MRI scanner toautomatically probe and adapt to its environment to prepare foroperation. Network connectivity allows the MRI system to be operatedfrom a mobile device such as a tablet, notepad or smart phone witheasy-to-use interfaces configured to automatically run desired scanningprotocols. Acquired images are immediately transferred to a secure cloudserver for data sharing, telemedicine and/or deep learning.

Following below are more detailed descriptions of various conceptsrelated to, and embodiments of, lower cost, lower power and/or portablelow-field MRI. It should be appreciated that the embodiments describedherein may be implemented in any of numerous ways. Examples of specificimplementations are provided below for illustrative purposes only. Itshould be appreciated that the embodiments and the features/capabilitiesprovided may be used individually, all together, or in any combinationof two or more, as aspects of the technology described herein are notlimited in this respect.

A significant contributor to the high cost, size, weight and powerconsumption of high-field MRI is the B₀ magnet itself along with theapparatus required to power the B₀ magnet and to perform thermalmanagement thereof. In particular, to produce the field strengthscharacteristic of high-field MRI, the B₀ magnet is typically implementedas an electromagnet configured in a solenoid geometry usingsuperconducting wires that need a cryogenic cooling system to keep thewires in a superconducting state. Not only is the superconductingmaterial itself expensive, but the cryogenic equipment to maintain thesuperconducting state is also expensive and complex.

The inventors have recognized that the low-field context allows for B₀magnet designs not feasible in the high-field regime. For example, dueat least in part to the lower field strengths, superconducting materialand the corresponding cryogenic cooling systems can be eliminated. Duein part to the low-field strengths, B₀ electromagnets constructed usingnon-superconducting material (e.g., copper) may be employed in thelow-field regime. However, such electromagnets still may consumerelatively large amounts of power during operation. For example,operating an electromagnet using a copper conductor to generate amagnetic field of 0.2 T or more requires a dedicated or specializedpower connection (e.g., a dedicated three-phase power connection). Theinventors have developed MRI systems that can be operated using mainselectricity (i.e., standard wall power), allowing the MRI system to bepowered at any location having common power connection, such as astandard wall outlet (e.g., 120V/20 A connection in the U.S.) or commonlarge appliance outlets (e.g., 220-240V/30 A). Thus, a low-power MRIsystem facilitates portability and availability, allowing an MRI systemto be operated at locations where it is needed (e.g., the MRI system canbe brought to the patient instead of vice versa), examples of which arediscussed in further detail below. In addition, operating from standardwall power eliminates the electronics conventionally needed to convertthree-phase power to single-phase power and to smooth out the powerprovided directly from the grid. Instead, wall power can be directlyconverted to DC and distributed to power the components of the MRIsystem.

FIGS. 2A and 2B illustrate a B₀ magnet formed using an electromagnet anda ferromagnetic yoke. In particular, B₀ magnet 200 is formed in part byan electromagnet 210 arranged in a bi-planar geometry comprisingelectromagnetic coils 212 a and 212 b on an upper side andelectromagnetic coils 214 a and 214 b on a lower side of B₀ magnet 200.According to some embodiments, the coils forming electromagnet 210 maybe formed from a number of turns of a copper wire or copper ribbon, orany other conductive material suitable for producing a magnetic fieldwhen operated (e.g., when electrical current is driven through theconductor windings). While the exemplary electromagnet illustrated inFIGS. 2A and 2B comprises two pairs of coils, an electromagnet may beformed using any number of coils in any configuration, as the aspectsare not limited in this respect. The electromagnetic coils formingelectromagnet 210 may be formed, for example, by winding a conductor 213(e.g., a copper ribbon, wire, paint, etc.) about a fiberglass ring 217.For example, conductor 213 may be a suitable insulated copper wire, oralternatively, conductor 213 may be a copper ribbon wound in conjunctionwith an insulating layer (e.g., a Mylar layer) to electrically isolatethe multiple windings of the coil. A connector 219 may be provided toallow for a power connection to provide current to operate coils 214 aand 214 b in series. A similar connector on the upper side of theelectromagnet (not visible in FIGS. 2A and 2B) may be provided tooperate coils 212 a and 212 b.

It should be appreciated that the electromagnetic coils may be formedfrom any suitable material and dimensioned in any suitable way so as toproduce or contribute to a desired B₀ magnetic field, as the aspects arenot limited for use with any particular type of electromagnet. As onenon-limiting example that may be suitable to form, in part, anelectromagnet (e.g., electromagnet 210), an electromagnetic coil may beconstructed using copper ribbon and mylar insulator having 155 turns toform an inner diameter of approximately 23-27 inches (e.g.,approximately 25 inches), an outer diameter of approximately 30-35inches (e.g., 32 inches). However, different material and/or differentdimensions may be used to construct an electromagnetic coil havingdesired characteristics, as the aspects are not limited in this respect.The upper and lower coil(s) may be positioned to provide a distance ofapproximately 10-15 inches (e.g., approximately 12.5 inches) between thelower coil on the upper side and the upper coil on the lower side. Itshould be appreciated that the dimensions will differ depending on thedesired characteristics including, for example, field strength, field ofview, etc.

In the exemplary B₀ magnet illustrated in FIGS. 2A and 2B, each coilpair 212 and 214 is separated by thermal management components 230 a and230 b, respectively, to transfer heat produced by the electromagneticcoils and gradient coils (not illustrated in FIGS. 2A and 2B) away fromthe magnets to provide thermal management for the MRI device. Inparticular, thermal management components 230 a and 230 b may comprise acooling plate having conduits that allow coolant to be circulatedthrough the cooling plate to transfer heat away from the magnets. Thecooling plate 230 a, 230 b may be constructed to reduce or eliminateeddy currents induced by operating the gradient coils that can produceelectromagnetic fields that disrupt the B₀ magnetic field produced bythe B0 magnet 200. For example, thermal management components 230 a and230 b may be the same or similar to any of the thermal managementcomponents described in U.S. application Ser. No. 14/846,042 entitled“Thermal Management Methods and Apparatus,” filed on Sep. 4, 2015, whichis incorporated by reference herein in its entirety. According to someembodiments, thermal management components may be eliminated, asdiscussed in further detail below.

B₀ magnet 200 further comprises a yoke 220 that is magnetically coupledto the electromagnet to capture magnetic flux that, in the absence ofyoke 220, would be lost and not contribute to the flux density in theregion of interest between the upper and lower electromagnetic coils. Inparticular, yoke 220 forms a “magnetic circuit” connecting the coils onthe upper and lower side of the electromagnet so as to increase the fluxdensity in the region between the coils, thus increasing the fieldstrength within the imaging region (also referred to as the field ofview) of the B₀ magnet. The imaging region or field of view defines thevolume in which the B₀ magnetic field produced by a given B0 magnet issuitable for imaging. More particularly, the imaging region or field ofview corresponds to the region for which the B₀ magnetic field issufficiently homogeneous at a desired field strength that detectable MRsignals are emitted by an object positioned therein in response toapplication of radio frequency excitation (e.g., a suitable radiofrequency pulse sequence). Yoke 220 comprises frame 222 and plates 224a, 224 b, which may be formed using any suitable ferromagnetic material(e.g., iron, steel, etc.). Plates 224 a, 224 b collect magnetic fluxgenerated by the coil pairs of electromagnet 210 and directs it to frame222 which, in turn, returns the flux back to the opposing coil pair,thereby increasing, by up to a factor of two, the magnetic flux densityin the imaging region between the coil pairs (e.g., coil pair 212 a, 212b and coil pair 214 a, 214 b) for the same amount of operating currentprovided to the coils. Thus, yoke 220 can be used to produce a higher B₀field (resulting in higher SNR) without a corresponding increase inpower requirements, or yoke 220 can be used to lower the powerrequirements of B₀ magnet 200 for a given B₀ field.

According to some embodiments, the material used for portions of yoke220 (i.e., frame 222 and/or plates 224 a, 224 b) is steel, for example,a low-carbon steel, silicon steel, cobalt steel, etc. According to someembodiments, gradient coils (not shown in FIGS. 2A, 2B) of the MRIsystem are arranged in relatively close proximity to plates 224 a, 224 binducing eddy currents in the plates. To mitigate, plates 224 a, 224 band/or frame 222 may be constructed of silicon steel, which is generallymore resistant to eddy current production than, for example, low-carbonsteel. It should be appreciated that yoke 220 may be constructed usingany ferromagnetic material with sufficient magnetic permeability and theindividual parts (e.g., frame 222 and plates 224 a, 224 b) may beconstructed of the same or different ferromagnetic material, as thetechniques of increasing flux density is not limited for use with anyparticular type of material or combination of materials. Furthermore, itshould be appreciated that yoke 220 can be formed using differentgeometries and arrangements.

It should be appreciated that the yoke 220 may be made of any suitablematerial and may be dimensioned to provide desired magnetic flux capturewhile satisfying other design constraints such as weight, cost, magneticproperties, etc. As an example, the frame of the yoke (e.g., frame 222)may be formed of a low-carbon steel of less than 0.2% carbon or siliconsteel, with the long beam(s) having a length of approximately 38 inches,a width of approximately 8 inches, and a thickness (depth) ofapproximately 2 inches, and the short beam(s) having a length ofapproximately 19 inches, a width of approximately 8 inches and athickness (depth of approximately 2 inches. The plates (e.g., plates 224a and 224 b) may be formed from a low-carbon steel of less than 0.2%carbon or silicon steel and have a diameter of approximately 30-35inches (e.g., approximately 32 inches). However, the above provideddimensions and materials are merely exemplary of a suitable embodimentof a yoke that can be used to capture magnetic flux generated by anelectromagnet.

As an example of the improvement achieved via the use of yoke 220,operating electromagnet 210 to produce a B₀ magnetic field ofapproximately 20 mT without yoke 220 consumes about 5 kW, whileproducing the same 20 mT B₀ magnetic field with yoke 220 consumes about750 W of power. Operating electromagnet 210 with the yoke 220, a B₀magnetic field of approximately 40 mT may be produced using 2 kW ofpower and a B₀ magnetic field of approximately 50 mT may be producedusing approximately 3 kW of power. Thus, the power requirements can besignificantly reduced by use of yoke 220 allowing for operation of a B₀magnet without a dedicated three-phase power connection. For example,mains electrical power in the United States and most of North America isprovided at 120V and 60 Hz and rated at 15 or 20 amps, permittingutilization for devices operating below 1800 and 2400 W, respectively.Many facilities also have 220-240 VAC outlets with 30 amp ratings,permitting devices operating up to 7200 W to be powered from suchoutlets. According to some embodiments, a low-field MRI system utilizinga B₀ magnet comprising an electromagnet and a yoke (e.g., B₀ magnet 200)is configured to be powered via a standard wall outlet, as discussed infurther detail below. According to some embodiments, a low-field MRIsystem utilizing a B₀ magnet comprising an electromagnet and a yoke(e.g., B₀ magnet 200) is configured to be powered via a 220-240 VACoutlet, as also discussed in further detail below.

Referring again to FIGS. 2A and 2B, exemplary B₀ magnet 210 furthercomprises shim rings 240 a, 240 b and shim disks 242 a, 242 b configuredto augment the generated B₀ magnetic field to improve homogeneity in thefield of view (e.g., in the region between the upper and lower coils ofthe electromagnet where the B₀ field is suitable for sufficient MRsignal production), as best seen in FIG. 2B in which the lower coilshave been removed. In particular, shim rings 240 and shim disk 242 aredimensioned and arranged to increase the uniformity of the magneticfield generated by the electromagnet at least within the field of viewof the B₀ magnet. In particular, the height, thickness and material ofshim rings 240 a, 240 b and the diameter, thickness and material of shimdisks 242 a, 242 b may be chosen so as to achieve a B₀ field of suitablehomogeneity. For example, the shim disk may be provided with a diameterof approximately 5-6 inches and a width of approximately 0.3-0.4 inches.A shim ring may be formed from a plurality of circular arc segments(e.g., 8 circular arc segments) each having a height of approximately20-22 inches, and a width of approximately 2 inches to form a ringhaving an inner diameter of approximately between 21-22 inches andapproximately between 23-24 inches.

The weight of the B₀ magnet is a significant portion of the overallweight of the MRI system which, in turn, impacts the portability of theMRI system. In embodiments that primarily use low carbon and/or siliconsteel for the yoke and shimming components, an exemplary B₀ magnet 200dimensioned similar to that described in the foregoing may weighapproximately 550 kilograms. According to some embodiments, cobalt steel(CoFe) may be used as the primary material for the yoke (and possiblythe shim components), potentially reducing the weight of B₀ magnet 200to approximately 450 Kilograms. However, CoFe is generally moreexpensive than, for example, low carbon steel, driving up the cost ofthe system. Accordingly, in some embodiments, select components may beformed using CoFe to balance the tradeoff between cost and weightarising from its use. Using such exemplary B₀ magnets a portable,cartable or otherwise transportable MRI system may be constructed, forexample, by integrating the B₀ magnet within a housing, frame or otherbody to which castors, wheels or other means of locomotion can beattached to allow the MRI system to be transported to desired locations(e.g., by manually pushing the MRI system and/or including motorizedassistance). As a result, an MRI system can be brought to the locationin which it is needed, increasing its availability and use as a clinicalinstrument and making available MRI applications that were previouslynot possible. According to some embodiments, the total weight of aportable MRI system is less than 1,500 pounds and, preferably, less than1000 pounds to facilitate maneuverability of the MRI system.

The primary contributor to the overall power consumption of a low-fieldMRI system employing a B₀ magnet such as B₀ magnet 200 is theelectromagnet (e.g., electromagnet 210). For example, in someembodiments, the electromagnet may consume 80% or more of the power ofthe overall MRI system. To significantly reduce the power requirementsof the MRI system, the inventors have developed B₀ magnets that utilizepermanent magnets to produce and/or contribute to the B₀ electromagneticfield. According to some embodiments, B₀ electromagnets are replacedwith permanent magnets as the main source of the B₀ electromagneticfield. A permanent magnet refers to any object or material thatmaintains its own persistent magnetic field once magnetized. Materialsthat can be magnetized to produce a permanent magnet are referred toherein as ferromagnetic and include, as non-limiting examples, iron,nickel, cobalt, neodymium (NdFeB) alloys, samarium cobalt (SmCo) alloys,alnico (AlNiCo) alloys, strontium ferrite, barium ferrite, etc.Permanent magnet material (e.g., magnetizable material that has beendriven to saturation by a magnetizing field) retains its magnetic fieldwhen the driving field is removed. The amount of magnetization retainedby a particular material is referred to as the material's remanence.Thus, once magnetized, a permanent magnet generates a magnetic fieldcorresponding to its remanence, eliminating the need for a power sourceto produce the magnetic field.

FIG. 3A illustrates a permanent B₀ magnet, in accordance with someembodiments. In particular, B₀ magnet 300 is formed by permanent magnets310 a and 310 b arranged in a bi-planar geometry and a yoke 320 thatcaptures electromagnetic flux produced by the permanent magnets andtransfers the flux to the opposing permanent magnet to increase the fluxdensity between permanent magnets 310 a and 310 b. Each of permanentmagnets 310 a and 310 b are formed from a plurality of concentricpermanent magnets. In particular, as visible in FIG. 3, permanentmagnetic 310 b comprises an outer ring of permanent magnets 314 a, amiddle ring of permanent magnets 314 b, an inner ring of permanentmagnets 314 c, and a permanent magnet disk 314 d at the center.Permanent magnet 310 a may comprise the same set of permanent magnetelements as permanent magnet 310 b.

The permanent magnet material used may be selected depending on thedesign requirements of the system. For example, according to someembodiments, the permanent magnets (or some portion thereof) may be madeof NdFeB, which produces a magnetic field with a relatively highmagnetic field per unit volume of material once magnetized. According tosome embodiments, SmCo material is used to form the permanent magnets,or some portion thereof. While NdFeB produces higher field strengths(and in general is less expensive than SmCo), SmCo exhibits less thermaldrift and thus provides a more stable magnetic field in the face oftemperature fluctuations. Other types of permanent magnet material(s)may be used as well, as the aspects are not limited in this respect. Ingeneral, the type or types of permanent magnet material utilized willdepend, at least in part, on the field strength, temperature stability,weight, cost and/or ease of use requirements of a given B₀ magnetimplementation.

The permanent magnet rings are sized and arranged to produce ahomogenous field of a desired strength in the central region (field ofview) between permanent magnets 310 a and 310 b. In the exemplaryembodiment illustrated in FIG. 3A, each permanent magnet ring comprisesa plurality segments, each segment formed using a plurality of blocksthat are stacked in the radial direction and positioned adjacent to oneanother about the periphery to form the respective ring. The inventorshave appreciated that by varying the width (in the direction tangent tothe ring) of each permanent magnet, less waste of useful space may beachieved while using less material. For example, the space betweenstacks that does not produce useful magnetic fields can be reduced byvarying the width of the blocks, for example, as function of the radialposition of the block, allowing for a closer fit to reduce wasted spaceand maximize the amount of magnetic field that can be generated in agiven space. The dimensions of the blocks may also be varied in anydesired way to facilitate the production of a magnetic field of desiredstrength and homogeneity, as discussed in further detail below.

B₀ magnet 300 further comprises yoke 320 configured and arranged tocapture magnetic flux generated by permanent magnets 310 a and 310 b anddirect it to the opposing side of the B₀ magnet to increase the fluxdensity in between permanent magnets 310 a and 310 b, increasing thefield strength within the field of view of the B₀ magnet. By capturingmagnetic flux and directing it to the region between permanent magnets310 a and 310 b, less permanent magnet material can be used to achieve adesired field strength, thus reducing the size, weight and cost of theB₀ magnet. Alternatively, for given permanent magnets, the fieldstrength can be increased, thus improving the SNR of the system withouthaving to use increased amounts of permanent magnet material. Forexemplary B₀ magnet 300, yoke 320 comprises a frame 322 and plates 324 aand 324 b. In a manner similar to that described above in connectionwith yoke 220, plates 324 a and 324 b capture magnetic flux generated bypermanent magnets 310 a and 310 b and direct it to frame 322 to becirculated via the magnetic return path of the yoke to increase the fluxdensity in the field of view of the B₀ magnet. Yoke 320 may beconstructed of any desired ferromagnetic material, for example, lowcarbon steel, CoFe and/or silicon steel, etc. to provide the desiredmagnetic properties for the yoke. According to some embodiments, plates324 a and 324 b (and/or frame 322 or portions thereof) may beconstructed of silicon steel or the like in areas where the gradientcoils could most prevalently induce eddy currents.

Exemplary frame 322 comprises arms 323 a and 323 b that attach to plates324 a and 324 b, respectively, and supports 325 a and 325 b providingthe magnetic return path for the flux generated by the permanentmagnets. The arms are generally designed to reduce the amount ofmaterial needed to support the permanent magnets while providingsufficient cross-section for the return path for the magnetic fluxgenerated by the permanent magnets. Arms 323 a has two supports within amagnetic return path for the B₀ field produced by the B₀ magnet.Supports 325 a and 325 b are produced with a gap 327 formed between,providing a measure of stability to the frame and/or lightness to thestructure while providing sufficient cross-section for the magnetic fluxgenerated by the permanent magnets. For example, the cross-sectionneeded for the return path of the magnetic flux can be divided betweenthe two support structures, thus providing a sufficient return pathwhile increasing the structural integrity of the frame. It should beappreciated that additional supports may be added to the structure, asthe technique is not limited for use with only two supports and anyparticular number of multiple support structures.

As discussed above, exemplary permanent magnets 310 a and 310 b comprisea plurality of rings of permanent magnetic material concentricallyarranged with a permanent magnet disk at the center. Each ring maycomprise a plurality of stacks of ferromagnetic material to form therespective ring, and each stack may include one or more blocks, whichmay have any number (including a single block in some embodiments and/orin some of the rings). The blocks forming each ring may be dimensionedand arranged to produce a desired magnetic field. The inventors haverecognized that the blocks may be dimensioned in a number of ways todecrease cost, reduce weight and/or improve the homogeneity of themagnetic field produced, as discussed in further detail in connectionwith the exemplary rings that together form permanent magnets of a B₀magnet, in accordance with some embodiments.

FIG. 3B illustrates a top-down view of a permanent magnet 310, whichmay, for example, be used as the design for permanent magnets 310 a and310 b of B₀ magnet 300 illustrated in FIG. 3A. Permanent magnet 310comprises concentric rings 310 a, 310 b, and 310 c, each constructed ofa plurality of stacks of ferromagnetic blocks, and a ferromagnetic disk310 d at the center. The direction of the frame of the yoke to whichpermanent magnet is attached is indicated by arrow 22. In embodiments inwhich the yoke is not symmetric (e.g., yoke 320), the yoke will causethe magnetic field produced by the permanent magnets for which itcaptures and focuses magnetic flux to be asymmetric as well, negativelyimpacting the uniformity of the B₀ magnetic field.

According to some embodiments, the block dimensions are varied tocompensate for the effects of the yoke on the magnetic field produced bythe permanent magnet. For example, dimensions of blocks in the fourregions 315 a, 315 b, 315 c and 315 d labeled in FIG. 3B may be varieddepending on which region the respective block is located. Inparticular, the height of the blocks (e.g., the dimension of the blocknormal to the plane of the circular magnet 310) may be greater in region315 c farthest away from the frame than corresponding blocks in region315 a closest to the frame. Block height can be varied in one or morerings or portions thereof, as the technique of compensating for theeffects of the yoke are not limited to varying any particular block, setof blocks and/or any particular dimension. One example of varying blockdimension to compensate for yoke effects are discussed in further detailbelow.

FIGS. 4A and 4B illustrate different views of an inner ring 410 (e.g.,ring 310 c illustrated in FIG. 3B), in accordance with some embodiments.Exemplary ring 410 includes a plurality (twelve in FIGS. 4A and 4B) ofstacks of two blocks each, thus forming two sub-rings of ferromagneticblocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring isformed of blocks (e.g., exemplary block 405 b) having a length x₀, awidth y₀ and a height (or depth) z₀. The outer sub-ring is formed ofblocks (e.g., exemplary block 405 a) having a length x₁, a width y₀ anda height (or depth) z₀. As shown, the blocks in the outer sub-ring havea greater length than block in the inner sub-ring (i.e., x₀<x₁),reducing the amount of empty space between adjacent block than if theblocks in the outer sub-ring were formed with a length x₀. As such moreof the space in which exemplary ring 410 is contained is occupied byfield producing magnetic material, increasing the field strength in thesame amount of space. It should be appreciated that the arrangement inexemplary ring 410 is merely illustrative and other arrangements ofblocks (e.g., number of stacks and number of blocks within each stack)may be used, as the aspects are not limited in this respect.

FIGS. 5A-C and 6A-C illustrate exemplary dimensions for blockscomprising the inner sub-ring and outer sub-ring of an inner permanentmagnet ring (e.g., exemplary dimensions for blocks 405 a and 405 bforming permanent magnet ring 410). In particular, exemplary block 505illustrated in FIG. 5A (e.g., a block in the inner sub-ring of innerpermanent magnet ring 410 or 310 c) may be manufactured to havedimensions x₀, y₀ and z₀. According to some embodiments, x₀ hasdimensions in a range between 20 and 25 millimeters, y₀ has dimensionsbetween 8 and 12 millimeter, and z₀ has dimensions between 19 and 23millimeters. Exemplary block 605 illustrated in FIG. 6A (e.g., a blockin the outer sub-ring of inner permanent magnet ring 410) may bemanufactured to have dimensions x₁, y₀ and z₀. According to someembodiments, x₁ has dimensions in a range between 27 and 32 millimeters.It should be appreciated that the dimensions of exemplary blocks 505 and605 are merely illustrative and the dimensions may be selected asdesired and are not limited in this respect. Additionally, blocks may beformed using any one or combination of ferromagnetic material, as theaspects are not limited for use with any particular type of magneticmaterial.

FIGS. 7A and 7B illustrate different views of a portion 715 of middlering 710 in a quadrant away from the yoke frame (e.g., the portion ofring 310 b in quadrant 315 c illustrated in FIG. 3B), in accordance withsome embodiments. Exemplary portion 715 of ring 710 includes a plurality(five in FIGS. 7A and 7B) of stacks of three blocks each, thus formingthree sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB,SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplaryblock 705 c) having a length x₂, a width y₁ and a height (or depth) z₁.The middle sub-ring is formed of blocks (e.g., exemplary block 705 b)having a length x₃, a width y₁ and a height (or depth) z₁. The outersub-ring is formed of blocks (e.g., exemplary block 705 a) having alength x₄, a width y₁ and a height (or depth) z₁. As shown, the blocksin the outer sub-ring have a greater length than blocks in the middlesub-ring which, in turn, have a length greater than the blocks in theinner sub-ring (i.e., x₂<x₃<x₄), reducing the amount of empty spacebetween adjacent block than if the blocks in all sub-rings were formedwith a length x₂. As such more of the space in which exemplary ring 710is contained is occupied by field producing magnetic material,increasing the field strength in the same amount of space. It should beappreciated that the arrangement in exemplary ring 710 is merelyillustrative and other arrangements of blocks (e.g., number of stacksand number of blocks within each stack) may be used, as the aspects arenot limited in this respect.

FIGS. 7C and 7D illustrate different views of a portion 715′ of middlering 710 in quadrant(s) in the middle with respect to the yoke frame(e.g., the portion of ring 310 b in quadrant 315 b and/or 315 dillustrated in FIG. 3B), in accordance with some embodiments. That is,portion 715′ may be used for both middle quadrants, for example, inembodiments where the middle quadrants are equidistant from the yokeframe. Exemplary portion 715′ of ring 710 includes a plurality (five inFIGS. 7C and 7D) of stacks of three blocks each, thus forming threesub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo,etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 705c′) having a length x₂, a width y₁ and a height (or depth) z₂. Themiddle sub-ring is formed of blocks (e.g., exemplary block 705 b′)having a length x₃, a width y₁ and a height (or depth) z₂. The outersub-ring is formed of blocks (e.g., exemplary block 705 a′) having alength x₄, a width y₁ and a height (or depth) z₂. As shown, the blocksin the outer sub-ring have a greater length than blocks in the middlesub-ring which, in turn, have a length greater than the blocks in theinner sub-ring (i.e., x₂<x₃<x₄), reducing the amount of empty spacebetween adjacent blocks than if the blocks in all sub-rings were formedwith a length x₂. As such more of the space in which exemplary ring 710is contained is occupied by field producing magnetic material,increasing the field strength in the same amount of space. It should beappreciated that the arrangement in exemplary ring 710 is merelyillustrative and other arrangements of blocks (e.g., number of stacksand number of blocks within each stack) may be used, as the aspects arenot limited in this respect.

FIGS. 7E and 7F illustrate different views of a portion 715″ of middlering 710 in a quadrant nearest the yoke frame (e.g., the portion of ring310 b in quadrant 315 a illustrated in FIG. 3B), in accordance with someembodiments. Exemplary portion 715″ of ring 710 includes a plurality(five in FIGS. 7E and 7F) of stacks of three blocks each, thus formingthree sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB,SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplaryblock 705 c″) having a length x₂, a width y₁ and a height (or depth) z₃.The middle sub-ring is formed of blocks (e.g., exemplary block 705 b″)having a length x₃, a width y₁ and a height (or depth) z₃. The outersub-ring is formed of blocks (e.g., exemplary block 705 a″) having alength x₄, a width y₁ and a height (or depth) z₃. As shown, the blocksin the outer sub-ring have a greater length than blocks in the middlesub-ring which, in turn, have a length greater than the blocks in theinner sub-ring (i.e., x₂<x₃<x₄), reducing the amount of empty spacebetween adjacent block than if the blocks in all sub-rings were formedwith a length x₂. As such more of the space in which exemplary ring 710is contained is occupied by field producing magnetic material,increasing the field strength in the same amount of space. It should beappreciated that the arrangement in exemplary ring 710 is merelyillustrative and other arrangements of blocks (e.g., number of stacksand number of blocks within each stack) may be used, as the aspects arenot limited in this respect.

FIGS. 8A-C, 9A-C and 10A-C illustrate exemplary dimensions for blockscomprising the inner, middle and outer sub-rings of a middle permanentmagnet ring (e.g., exemplary dimensions for blocks 705 a-705 c, 705a′-705 c′ and 705 a″-705 c″ forming permanent magnet ring 710illustrated in FIGS. 7A-7F). In particular, exemplary block 805illustrated in FIG. 8A (e.g., a block in the inner sub-ring of middlepermanent magnet ring 710 or 310 b) may be manufactured to havedimensions x₂, y₁ and z_(n) as labeled in FIGS. 8B and 8C. According tosome embodiments, x₂ has dimensions in a range between 31 and 35millimeters, y₁ has dimensions between 6 and 10 millimeters, and z_(n)has dimensions between 21 and 25 millimeters. Exemplary block 905illustrated in FIG. 9A (e.g., a block in the middle sub-ring of middlepermanent magnet ring 710 or 310 b) may be manufactured to havedimensions x₃, y₁ and z_(n) as labeled in FIGS. 9B and 9C. According tosome embodiments, x₃ has dimensions in a range between 34 and 38millimeters. Similarly, exemplary block 1005 illustrated in FIG. 10A(e.g., a block in the outer sub-ring of middle permanent magnet ring 710or 310 b) may be manufactured to have dimensions x₄, y₁ and z_(n) aslabeled in FIGS. 10B and 10C. According to some embodiments, x₄ hasdimensions in a range between 37 and 41 millimeters. It should beappreciated that the dimensions of exemplary blocks 805, 905 and 1005are merely illustrative and the dimensions may be selected as desiredand are not limited in this respect. Additionally, blocks may be formedusing any one or combination of ferromagnetic material, as the aspectsare not limited for use with any particular type of magnetic material.As discussed above, the height of the blocks may be varied to compensatefor effects in the homogeneity of the magnetic field resulting from thepresence of the yoke. According to some embodiments, z_(n) is varieddepending on which quadrant the block appears in (e.g., whether theblock is in quadrant 715, 715′ or 715″), further details of which arediscussed below.

FIGS. 11A and 11B illustrate different views of a portion 1115 of outerring 1110 in a quadrant away from the yoke frame (e.g., the portion ofring 310 a in quadrant 315 c illustrated in FIG. 3B), in accordance withsome embodiments. Exemplary portion 1115 of ring 1110 includes aplurality (nine in FIGS. 11A and 11B) of stacks of three blocks each,thus forming three sub-rings of ferromagnetic blocks (e.g., blocksformed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks(e.g., exemplary block 1105 c) having a length x₅, a width y₂ and aheight (or depth) z₄. The middle sub-ring is formed of blocks (e.g.,exemplary block 1105 b) having a length x₆, a width y₂ and a height (ordepth) z₄. The outer sub-ring is formed of blocks (e.g., exemplary block1105 a) having a length x₇, a width y₂ and a height (or depth) z₄. Asshown, the blocks in the outer sub-ring have a greater length thanblocks in the middle sub-ring which, in turn, have a length greater thanthe blocks in the inner sub-ring (i.e., x₅<x₆<x₇), reducing the amountof empty space between adjacent block than if the blocks in allsub-rings were formed with a length x₅. As such more of the space inwhich exemplary ring 1110 is contained is occupied by field producingmagnetic material, increasing the field strength in the same amount ofspace. It should be appreciated that the arrangement in exemplary ring1110 is merely illustrative and other arrangements of blocks (e.g.,number of stacks and number of blocks within each stack) may be used, asthe aspects are not limited in this respect.

FIGS. 11C and 11D illustrate different views of a portion 1115′ of outerring 1110 in quadrant(s) in the middle with respect to the yoke frame(e.g., the portion of ring 310 a in quadrant 315 b and/or 315 dillustrated in FIG. 3B), in accordance with some embodiments. That is,portion 1115′ may be used for both middle quadrants, for example, inembodiments where the middle quadrants are equidistant from the yokeframe. Exemplary portion 1115′ of ring 1110 includes a plurality (ninein FIGS. 11C and 11D) of stacks of three blocks each, thus forming threesub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo,etc.). The inner sub-ring is formed of blocks (e.g., exemplary block1105 c′) having a length x₅, a width y₂ and a height (or depth) z₅. Themiddle sub-ring is formed of blocks (e.g., exemplary block 1105 b′)having a length x₆, a width y₂ and a height (or depth) z₅. The outersub-ring is formed of blocks (e.g., exemplary block 1105 a′) having alength x₇, a width y₂ and a height (or depth) z₅. As shown, the blocksin the outer sub-ring have a greater length than blocks in the middlesub-ring which, in turn, have a length greater than the blocks in theinner sub-ring (i.e., x₅<x₆<x₇), reducing the amount of empty spacebetween adjacent block than if the blocks in all sub-rings were formedwith a length x₅. As such more of the space in which exemplary ring 1110is contained is occupied by field producing magnetic material,increasing the field strength in the same amount of space. It should beappreciated that the arrangement in exemplary ring 1110 is merelyillustrative and other arrangements of blocks (e.g., number of stacksand number of blocks within each stack) may be used, as the aspects arenot limited in this respect.

FIGS. 11E and 11F illustrate different views of a portion 1115″ of outerring 1110 in a quadrant nearest the yoke frame (e.g., the portion ofring 310 a in quadrant 315 a illustrated in FIG. 3B), in accordance withsome embodiments. Exemplary portion 1115″ of ring 1110 includes aplurality (nine in FIGS. 11E and 11F) of stacks of three blocks each,thus forming three sub-rings of ferromagnetic blocks (e.g., blocksformed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks(e.g., exemplary block 1105 c″) having a length x₅, a width y₂ and aheight (or depth) z₆. The middle sub-ring is formed of blocks (e.g.,exemplary block 1105 b″) having a length x₆, a width y₂ and a height (ordepth) z₆. The outer sub-ring is formed of blocks (e.g., exemplary block1105 a″) having a length x₇, a width y₂ and a height (or depth) z₆. Asshown, the blocks in the outer sub-ring have a greater length thanblocks in the middle sub-ring which, in turn, have a length greater thanthe blocks in the inner sub-ring (i.e., x₅<x₆<x₇), reducing the amountof empty space between adjacent block than if the blocks in allsub-rings were formed with a length x₅. As such more of the space inwhich exemplary ring 1110 is contained is occupied by field producingmagnetic material, increasing the field strength in the same amount ofspace. It should be appreciated that the arrangement in exemplary ring1110 is merely illustrative and other arrangements of blocks (e.g.,number of stacks and number of blocks within each stack) may be used, asthe aspects are not limited in this respect.

FIGS. 12A-C, 13A-C and 14A-C illustrate exemplary dimensions for blockscomprising the inner, middle and outer sub-rings of an outer permanentmagnet ring (e.g., exemplary dimensions for blocks 1105 a-1105 c, 1105a′-1105 c′ and 1105 a″-1105 c″ forming permanent magnet ring 1110). Inparticular, exemplary block 1205 illustrated in FIG. 12A (e.g., a blockin the inner sub-ring of outer permanent magnet ring 1110 or 310 a) maybe manufactured to have dimensions x₅, y₂ and z_(i), as labeled in FIGS.12B and 12C. According to some embodiments, x₅ is in a range between 34and 38 millimeters, y₂ is in a range between 16 and 20 millimeters andz_(i) is in a range between 22 and 27 millimeters. Exemplary block 1305illustrated in FIG. 13A (e.g., a block in the middle sub-ring of outerpermanent magnet ring 1110 or 310 a) may be manufactured to havedimensions x₆, y₂ and z_(i), as labeled in FIGS. 13B and 13C. Accordingto some embodiments, x₆ is in a range between 37 and 43 millimeters, y₂is in a range between 16 and 20 millimeters and z_(i) is in a rangebetween 22 and 27 millimeters. Similarly, exemplary block 1405illustrated in FIG. 14A (e.g., a block in the outer sub-ring of outerpermanent magnet ring 1110 or 310 a) may be manufactured to havedimensions x₇, y₂ and z_(i), as labeled in FIGS. 14B and 14C. Accordingto some embodiments, x₇ is in a range between 40 and 45 millimeters, y₂is in a range between 16 and 20 millimeters and z_(i) is in a rangebetween 22 and 27 millimeters. It should be appreciated that thedimensions of exemplary blocks 1205, 1305 and 1405 are merelyillustrative and the dimensions may be selected as desired and are notlimited in this respect. Additionally, blocks may be formed using anyone or combination of ferromagnetic material, as the aspects are notlimited for use with any particular type of magnetic material. Asdiscussed above, the height of the blocks may be varied to compensatefor effects in the homogeneity of the magnetic field resulting from thepresence of the yoke. According to some embodiments, z_(i) is varieddepending on which quadrant the block appears in (e.g., whether theblock is in quadrant 1115, 1115′ or 1115″), further details of which arediscussed below.

It should be appreciated that the permanent magnet illustrated in FIG.3A can be manufactured using any number and arrangement of permanentmagnet blocks and are limited to the number, arrangement, dimensions ormaterials illustrated herein. The configuration of the permanent magnetswill depend, at least in part, on the design characteristics of the B₀magnet, including, but not limited to, the field strength, field ofview, portability and/or cost desired for the MRI system in which the B₀magnet is intended to operate. For example, the permanent magnet blocksmay be dimensioned to produce a magnetic field ranging from 20 mT to 0.1T, depending on the field strength desired. However, it should beappreciated that other low-field strengths (e.g., up to approximately0.2 T) may be produced by increasing the dimensions of the permanentmagnet, though such increases will also increase the size, weight andcost of the B₀ magnet.

As discussed above, the height or depth of the blocks used in thedifferent quadrants may be varied to compensate for effects on the B₀magnetic field resulting from an asymmetric yoke. For example, in theconfiguration illustrated in FIG. 3A, the position of frame 322 (inparticular, legs 325 a and 325 b) to the permanent magnets 310 a and 310b results in magnetic flux being drawn away from regions proximate theframe (e.g., quadrant 315 a), reducing the flux density in theseregions. To address the resulting non-uniformity in the magnetic field,the height or depth of the blocks in affected regions may be varied(e.g., increased) to generate additional magnetic flux to compensate forthe reduction in magnetic flux density caused by the yoke, therebyimproving the homogeneity of the B₀ magnetic field within the field ofview of the B₀ magnet.

The inventors have appreciated that the arrangement, dimensions andmaterials used for the permanent magnet blocks may be chosen to minimizethe Lorentz forces produced by the B₀ coil during operation of thegradient coils. This technique can be used to reduce vibration andacoustic noise during the operation of the MRI system. According to someembodiments, the design of the permanent magnet blocks are chosen toreduce magnetic field components perpendicular to the B₀ field, i.e.,parallel to the plane of the gradient coils. According to someembodiments, the outer ring of permanent magnet blocks are designed toreduce the magnetic field components responsible for vibration of thegradient coils during operation in areas outside the field of view ofthe MRI system, thereby reducing vibration and acoustic noise generatedduring operation of the MRI system.

FIGS. 15A-15C illustrate an exemplary permanent magnet disk (e.g.,permanent magnet disk 310 d illustrated in FIG. 3B), in accordance withsome embodiment. Permanent magnet disk 1510 is configured to be placedat center of the permanent magnet (e.g., permanent magnet 310 a and/or310 b illustrated in FIG. 3A) to contribute to the B₀ field produced bythe permanent magnet. Permanent magnet disk 1510 may be formed from anysuitable ferromagnetic material (e.g., NdFeB, SmCo, etc.) and havesuitable dimensions so that, when magnetized, permanent magnet diskproduces a desired magnetic field. Exemplary permanent magnet disk 1510has a diameter D (e.g., in range between 32 and 36 millimeters) and athickness z₇ (e.g., in range between 18 and 22 millimeters, though anydimensions may be used to satisfy the design requirements of aparticular B₀ magnet (e.g., to achieve a desired field strength and/orhomogeneity).

FIG. 16 illustrates a B₀ magnet 1600, in accordance with someembodiments. B₀ magnet 1600 may share design components with B₀ magnet300 illustrated in FIG. 3A. In particular, B₀ magnet 1600 is formed bypermanent magnets 1610 a and 1610 b arranged in a bi-planar geometrywith a yoke 1620 coupled thereto to capture electromagnetic fluxproduced by the permanent magnets and transfer the flux to the opposingpermanent magnet to increase the flux density between permanent magnets1610 a and 1610 b. Each of permanent magnets 1610 a and 1610 b areformed from a plurality of concentric permanent magnets, as shown bypermanent magnet 1610 b comprising an outer ring of permanent magnets1614 a, a middle ring of permanent magnets 1614 b, an inner ring ofpermanent magnets 1614 c, and a permanent magnet disk 1614 d at thecenter. Permanent magnet 1610 a may comprise the same set of permanentmagnet elements as permanent magnet 1610 b. The permanent magnetmaterial used may be selected depending on the design requirements ofthe system (e.g., NdFeB, SmCo, etc. depending on the propertiesdesired).

The permanent magnet rings are sized and arranged to produce ahomogenous field of a desired strength in the central region (field ofview) between permanent magnets 1610 a and 1610 b. In particular, in theexemplary embodiment illustrated in FIG. 16, each permanent magnet ringcomprises a plurality of circular arc segments sized and positioned toproduce a desired B₀ magnetic field, as discussed in further detailbelow. In a similar manner to yoke 320 illustrated in FIG. 3A, yoke 1620is configured and arranged to capture magnetic flux generated bypermanent magnets 1610 a and 1610 b and direct it to the opposing sideof the B₀ magnet to increase the flux density in between permanentmagnets 1610 a and 1610 b. Yoke 1620 thereby increases the fieldstrength within the field of view of the B₀ magnet with less permanentmagnet material, reducing the size, weight and cost of the B₀ magnet.Yoke 1620 also comprises a frame 1622 and plates 1624 a and 1624 b that,in a manner similar to that described above in connection with yoke1620, captures and circulates magnetic flux generated by the permanentmagnets 1610 a and via the magnetic return path of the yoke to increasethe flux density in the field of view of the B₀ magnet. The structure ofyoke 1620 may be similar to that described above to provide sufficientmaterial to accommodate the magnetic flux generated by the permanentmagnets and providing sufficient stability, while minimizing the amountof material used to, for example, reduce the cost and weight of the B₀magnet.

FIG. 17 illustrates a top-down view of a permanent magnet 1710, whichmay, for example, be used as the design for permanent magnets 1710 a and1710 b of B₀ magnet 1600 illustrated in FIG. 16. Permanent magnet 1710comprises concentric rings 1710 a, 1710 b, and 1710 c, each constructedof a plurality of circular arc segments of ferromagnetic material, and aferromagnetic disk 1710 d at the center. The direction of the frame ofthe yoke to which permanent magnet is attached is indicated by arrow 22.In embodiments in which the yoke is not symmetric (e.g., yoke 1620), theyoke will cause the magnetic field produced by the permanent magnets forwhich it captures and focuses magnetic flux to be asymmetric as well,negatively impacting the uniformity of the B₀ magnetic field. Accordingto some embodiments, one or more dimensions of the circular arc segmentsare varied to compensate for the effects of the yoke on the magneticfield produced by the permanent magnet. For example, one or moredimensions of circular arc segments in the four quadrants 1715 a, 1715b, 1715 c and 1715 d labeled in FIG. 17 may be varied to compensate forthe effects of the yoke on the B₀ magnetic field, as discussed infurther detail below.

FIGS. 18A and 18B illustrate different views of an inner ring 1810(e.g., ring 1710 c illustrated in FIG. 17), in accordance with someembodiments. Exemplary ring 1810 includes a plurality (eight inexemplary ring 1810 illustrated in FIGS. 18A and 18B) of ferromagneticcircular arc segments (e.g., segments formed of NdFeB, SmCo, etc.), eachspanning 45° of the ring. In exemplary ring 1810, the circular arcsegments (e.g., exemplary circular arc segment 1805) are dimensioned soas to provide a ring with inner radius R1 and outer radius R2 and aheight or depth z₈. According to some embodiments, inner ring 1810 hasdimensions of R1 between 45-47 mm (e.g., 46.08 mm), R2 between 62-64 mm(e.g., 62.91 mm) and z₈ between 22 and 25 mm (e.g., 23.46 mm). It shouldbe appreciated that the number of circular arc segments and thedimensions thereof may be chosen as desired to produce a desired B₀magnetic field (e.g., a desired field strength and/or homogeneity), asthe aspects are not limited in this respect.

FIGS. 18C and 18D illustrate different views of a segment 1815 that canbe used to form middle ring 1810 illustrated in FIG. 18E (e.g., ring1710 b illustrated in FIG. 17). For example, segment 1815 can be used toprovide the segments in quadrants Q1-Q4 as illustrated in FIG. 18E(also, e.g., segments in quadrants 1715 a-d of ring 1710 b illustratedin FIG. 17). Exemplary portion 1815′ includes a plurality offerromagnetic circular arc segments (e.g., segments formed of NdFeB,SmCo, etc.), In FIGS. 18C-18E, two circular arc segments (e.g.,exemplary circular arc segment 1805′), each spanning 45°, form aquadrant of ring 1810′. In exemplary portion 1815′ of ring 1810′, thecircular arc segments are dimensioned so as to provide a ring with innerradius R1 and outer radius R2 and a height or depth z₉, which dimensionscan be chosen for each quadrant to achieve a desired magnetic field,non-limiting examples of which are provided below.

FIGS. 18F and 18G illustrate different views of a segment 1815 that canbe used to form outer ring 1810″ illustrated in FIG. 18H (e.g., ring1710 a illustrated in FIG. 17). For example, segment 1815″ can be usedto provide the segments in quadrants Q1-Q4 as illustrated in FIG. 18H(also, e.g., segments in quadrants 1715 a-d of ring 1710 a illustratedin FIG. 17). Exemplary portion 1815″ includes a plurality offerromagnetic circular arc segments (e.g., segments formed of NdFeB,SmCo, etc.), In FIGS. 18F-18H, five circular arc segments (e.g.,exemplary circular arc segment 1805″), each spanning 18° of ring 1810″,form a quadrant of ring 1810″. In exemplary segment 1815 of ring 1810″,the circular arc segments are dimensioned so as to provide a ring withinner radius R1 and outer radius R2 and a height or depth z₁₀, whichdimensions can be chosen for each quadrant to achieve a desired magneticfield.

As discussed above, the inventors have developed low power, portablelow-field MRI systems that can be deployed in virtually any environmentand that can be brought to the patient who will undergo an imagingprocedure. In this way, patients in emergency rooms, intensive careunits, operating rooms and a host of other locations can benefit fromMRI in circumstances where MRI is conventionally unavailable. Aspectsthat facilitate portable MRI are discussed in further detail below.

FIGS. 19A and 19B illustrate a low power, portable low-field MRI system,in accordance with some embodiments. Portable MRI system 1900 comprisesa B₀ magnet 1905 including at least one first permanent magnet 1910 aand at least one second permanent magnet 1910 b magnetically coupled toone another by a ferromagnetic yoke 1920 configured to capture andchannel magnetic flux to increase the magnetic flux density within theimaging region (field of view) of the MRI system. Permanent magnets 1910a and 1910 b may be constructed using any suitable technique, includingany of the techniques described herein (e.g., using any of thetechniques, designs and/or materials described in connection with B₀magnet 300 illustrated in FIG. 3A and/or B₀ magnet 1600 illustrated inFIG. 16 and described in the accompanying description thereof). Yoke1920 may also be constructed using any of the techniques describedherein (e.g., using any of the techniques, designs and/or materialsdescribed in connection with yokes 320 and 1620 illustrated in FIG. 3Aand FIG. 16 and described in the accompanying description thereof). Itshould be appreciated that, in some embodiments, B₀ magnet 1905 may beformed using electromagnets using any of the electromagnet techniquesdescribed herein (e.g., using any of the techniques, designs and/ormaterials described in connection with B₀ magnet 200 illustrated inFIGS. 2A and 2B and described in the accompanying description thereof).B₀ magnet 1905 may be encased or enclosed in a housing 1912 along withone or more other magnetics components, such as the system's gradientcoils (e.g., x-gradient, y-gradient and z-gradient coils) and/or anyshim components (e.g., shim coils or permanent magnetic shims), B₀correction coils, etc.

B₀ magnet 1905 may be coupled to or otherwise attached or mounted tobase 1950 by a positioning mechanism 1990, such as a goniometric stage(examples of which are discussed in further detail below in connectionwith FIGS. 45A-D and 46A-B), so that the B₀ magnet can be tilted (e.g.,rotated about its center of mass) to provide an incline to accommodate apatient's anatomy as needed. In FIG. 19A, the B₀ magnet is shown levelwithout an incline and, in FIG. 19B, the B₀ magnet is shown afterundergoing a rotation to incline the surface supporting the patient'sanatomy being scanned. Positioning mechanism 1990 may be fixed to one ormore load bearing structures of base 1950 arranged to support the weightof B₀ magnet 1900.

In addition to providing the load bearing structures for supporting theB₀ magnet, base 1950 also includes an interior space configured to housethe electronics 1970 needed to operate the portable MRI system 1900. Forexample, base 1950 may house the power components to operate thegradient coils (e.g., X, Y and Z) and the RF transmit/receive coils. Theinventors have developed generally low power, low noise and low costgradient amplifiers configured to suitably power gradient coils in thelow-field regime, designed to be relatively low cost, and constructedfor mounting within the base of the portable MRI system (i.e., insteadof being statically racked in a separate room of a fixed installment asis conventionally done). Examples of suitable power components tooperate the gradient coils are described in further detail below (e.g.,the power components described in connection with FIGS. 20-34).According to some embodiments, the power electronics for powering thegradient coils of an MRI system consume less than 50 W when the systemis idle and between 100-300 W when the MRI system is operating (i.e.,during image acquisition). Base 1950 may also house the RF coilamplifiers (i.e., power amplifiers to operate the transmit/receive coilsof the system), power supplies, console, power distribution unit andother electronics needed to operate the MRI system, further details ofwhich are described below.

According to some embodiments, the electronics 1970 needed to operateportable MRI system 1900 consume less than 1 kW of power, in someembodiments, less than 750 W of power and, in some embodiments, lessthan 500 W of power (e.g., MRI systems utilizing a permanent B₀ magnetsolution). Techniques for facilitating low power operation of an MRIdevice are discussed in further detail below. However, systems thatconsume greater power may also be utilized as well, as the aspects arenot limited in this respect. Exemplary portable MRI system 1900illustrated in FIGS. 19A and 19B may be powered via a single powerconnection 1975 configured to connect to a source of mains electricity,such as an outlet providing single-phase power (e.g., a standard orlarge appliance outlet). Accordingly, the portable MRI system can beplugged into a single available power outlet and operated therefrom,eliminating the need for a dedicated power source (e.g., eliminating theneed for a dedicated three-phase power source as well as eliminating theneed for further power conversion electronics to convert three phasepower to single phase power to be distributed to correspondingcomponents of the MRI system) and increasing the availability of the MRIsystem and the circumstances and locations in which the portable MRIsystem may be used.

Portable MRI system 1900 illustrated in FIGS. 19A and 19B also comprisesa conveyance mechanism 1980 that allows the portable MRI system to betransported to different locations. The conveyance mechanism maycomprise one or more components configured to facilitate movement of theportable MRI system, for example, to a location at which MRI is needed.According to some embodiments, conveyance mechanism comprises a motor1986 coupled to drive wheels 1984. In this manner, conveyance mechanism1980 provides motorized assistance in transporting MRI system 1900 todesired locations. Conveyance mechanism 1980 may also include aplurality of castors 1982 to assist with support and stability as wellas facilitating transport.

According to some embodiments, conveyance mechanism 1980 includesmotorized assistance controlled using a controller (e.g., a joystick orother controller that can be manipulated by a person) to guide theportable MRI system during transportation to desired locations.According to some embodiments, the conveyance mechanism comprises powerassist means configured to detect when force is applied to the MRIsystem and to, in response, engage the conveyance mechanism to providemotorized assistance in the direction of the detected force. Forexample, rail 1955 of base 1950 illustrated in FIGS. 19A and 19B may beconfigured to detect when force is applied to the rail (e.g., bypersonnel pushing on the rail) and engage the conveyance mechanism toprovide motorized assistance to drive the wheels in the direction of theapplied force. As a result, a user can guide the portable MRI systemwith the assistance of the conveyance mechanism that responds to thedirection of force applied by the user. The power assist mechanism mayalso provide a safety mechanism for collisions. In particular, the forceof contact with another object (e.g., a wall, bed or other structure)may also be detected and the conveyance mechanism will react accordinglywith a motorized locomotion response away from the object. According tosome embodiments, motorized assistance may be eliminated and theportable MRI system may be transported by having personnel move thesystem to desired locations using manual force.

Portable MRI system 1900 includes slides 1960 that provideelectromagnetic shielding to the imaging region of the system. Slides1960 may be transparent or translucent to preserve the feeling ofopenness of the MRI system to assist patients who may experienceclaustrophobia during conventional MRI performed within a closed bore.Slides 1960 may also be perforated to allow air flow to increase thesense of openness and/or to dissipate acoustic noise generated by theMRI system during operation. The slides may have shielding 1965incorporated therein to block electromagnetic noise from reaching theimaging region. According to some embodiments, slides 1960 may also beformed by a conductive mesh providing shielding 1965 to the imagingregion and promoting a sense of openness for the system. Thus, slides1960 may provide electromagnetic shielding that is moveable to allow apatient to be positioned within the system, permitting adjustment bypersonnel once a patient is positioned or during acquisition, and/orenabling a surgeon to gain access to the patient, etc. Thus, themoveable shielding facilitates flexibility that allows the portable MRIsystem to not only be utilized in unshielded rooms, but enablesprocedures to be performed that are otherwise unavailable. Exemplaryslides providing varying levels of electromagnetic shielding arediscussed in further detail below.

According to some embodiments, a portable MRI system does not includeslides, providing for a substantially open imaging region, facilitatingeasier placement of a patient within the system, reducing the feeling ofclaustrophobia and/or improving access to the patient positioned withinthe MRI system (e.g., allowing a physician or surgeon to access thepatient before, during or after an imaging procedure without having toremove the patient from the system). The inventors have developedtechniques that facilitate performing MRI with varying levels ofelectromagnetic shielding, including no or substantially no shielding ofthe imaging region, including a noise suppression system adapted tosuppress electromagnetic noise in the environment. According to someembodiments, portable MRI system 1900 may be equipped with a noisereduction system using one or more of the noise suppression and/oravoidance techniques described herein to, for example, dynamically adaptthe noise suppression/cancellation response in concert with theshielding configuration of a given shielding arrangement of the portableMRI system 1900. Thus, portable low field MRI system 1900 can betransported to the patient and/or to a desired location and operatedoutside specially shielded rooms (e.g., in an emergency room, operatingroom, NICU, general practitioner's office, clinic) and/or broughtbedside directly to the patient wherever located, allowing for MRI to beperformed when and where it is needed. To facilitate portable MRI thatcan be operated in virtually any location, the inventors have developedlow power MRI systems that, in accordance with some embodiments, areconfigured to be powered by main electricity (e.g., single-phaseelectric power from standard or industrial wall outlets), as discussedin further detail below.

As discussed above, conventional MRI systems consume significant power,requiring dedicated three-phase power sources to operate. In particular,conventional MRI systems that use superconducting material to form theB₀ magnet require cryogenic cooling systems that consume substantialpower to keep the conductors in a superconducting state. In addition,the power amplifiers used to operate the gradient amplifiers are largepower components that draw large amounts of power and are typicallystored in a separate room that houses the electronic components of thesystem. Moreover, power components configured to operate thetransmit/receive coil systems of conventional MRI system also consumesignificant amounts of power. Many conventional high field MRI systemsare require HVAC systems that also draw substantial amounts of power.

Conventional MRI systems are fixed installments requiring a specializedand dedicated spaces. As a result, the requirement of a dedicatedthree-phase power connection to operate the MRI system is not a criticallimitation for these systems, as it is just one of a number of dedicatedand specialized features of a conventional MRI installment. However,requiring a dedicated three-phase power source places significantrestrictions on locations at which a portable MRI system can beoperated. Accordingly, the inventors have developed a low power MRIsystem that facilitates portability of the MRI system. For example, inaccordance with some embodiments, a low power MRI system is configuredto operate using mains power (e.g., single phase electric power from astandard or industrial outlet). Exemplary aspects of a low power MRIsystem are discussed in further detail below.

According to some embodiments, a low power MRI system comprises apermanent B₀ magnet (e.g., any of the permanent magnets discussed hereinsuch as those illustrated in FIGS. 3A and 16). Because a permanent B₀magnet, once magnetized, will produce its own persistent magnetic field,power is not required to operate the permanent B₀ magnet to generate itsmagnetic field. As a result, a significant (often dominant) contributorto the overall power consumption of an MRI system can be eliminated,facilitating the development of an MRI system that can be powered usingmains electricity (e.g., via a standard wall outlet or common largehousehold appliance outlets), as discussed in further detail below inconnection with exemplary low power MRI systems.

Furthermore, conventional power components adapted to operate a gradientcoil system are generally unsuitable for use in low-field MRI due, atleast in part to, expense and noise levels and are unsuitable for lowpower and/or portable MRI due to power consumption, size and weight. Forexample, while the cost of conventional power components used to operategradient coils in currently available MRI systems may be acceptablegiven the relative insignificance compared to the total cost of ahigh-field MRI installation, this cost may be unacceptably high in thecontext of a low-field MRI system that is designed as a lower costalternative. Thus, the cost of a power component conventionally used forhigh-field MRI may be disproportionately large and therefore notsatisfactory for some lower cost low-field MRI systems.

Additionally, the relatively low SNR in the low-field (and particularlyin the very-low and ultra-low-field regimes) renders conventionalgradient coil power components unsuitable. In particular, conventionalpower components for driving gradient coils are generally unsuitable forlow-field MRI systems because they are not designed to drive the coilswith sufficiently low noise. Although the noise injected by such powercomponents may be acceptable in the high SNR regime of high-field MRIsystems, such components generally do not provide a sufficiently lowlevel of noise to provide acceptable image quality in a low-field MRIsystem. For example, conventional power components may exhibitunsatisfactory variation in the output (e.g., ripple) for use in thelow-field context, injecting relatively significant noise into thegradient coil system of a low-field MRI system.

Moreover, conventional power components configured to drive the gradientcoil system of currently available MRI systems are not designed to bepower efficient, consuming large amounts of power. In addition,conventional power components configured to operate the gradient coilsystem of currently available MRI systems are large, heavy devices,typically racked in a separate room adjacent the MRI device along withthe other electronic components. Thus, conventional gradient powercomponents are not suitable for use in a low power, portable MRI system.

The inventors have developed low-noise, low power gradient powercomponent(s) suitable for driving the gradient coil system of alow-field MRI system. In particular, techniques developed by theinventors provide for a low cost, low power, low noise gradient coilsystem suitable for a low-field, very-low field or ultra-low field MRIsystem, and more particularly, for a portable MRI system that canoperate using standard and/or commonly available power connections. Thatis, in addition to facilitating a low power MRI system, the gradientcoils and gradient coil power components facilitate MRI at lower fieldstrengths not attainable using conventional gradient coil systems due,at least in part, to the low noise operation of the gradient powercomponents. According to some embodiments, the power electronics forpowering the gradient coils of an MRI system consume less than 50 W whenthe system is idle and between 100-300 W when the MRI system isoperating (i.e., during image acquisition), allowing for operation fromstandard wall power, some examples of which are described in furtherdetail below in connection with FIGS. 20-34.

FIG. 20 illustrates drive circuitry for driving a current through a coil2002 of an MRI system to produce a magnetic field in accordance with adesired pulse sequence, according to some embodiments. Power component1914 drives a current through coil 2002 based on a control signal fromthe controller 1906. The controller 1906 may produce a control signal todrive power component 1914 based on a pulse sequence implemented bycontroller 1906 (or provided by one or more other controllers), asdiscussed above. In some embodiments, coil 2002 may be a gradient coil1928. However, the techniques described herein are not limited in thisrespect, as coil 2002 may be a coil of magnet 1922, shim coil 1924, oran RF transmit and/or receive coil 1926.

Power components configured to power gradient coils typically providerelatively high power and typically need to provide precise control overthe current provided to the gradient coil so that the desired pulsesequence can be delivered faithfully. Inaccuracies in delivering thecommanded current to the gradient coil results in a decrease insignal-to-noise ratio due to differences between the gradient pulsesequence being delivered and the intended (and expected) pulse sequence.Power components configured to drive gradient coils also should to beresponsive in delivering the commanded current to the gradient coil,including rapid transition between commanded current levels so as tofaithfully produce the current waveforms required by the desired pulsesequences. Accordingly, the inventors have developed power componentscapable of being controlled to accurately and precisely provide current,with relatively low noise and relatively high efficiency, to one or moregradient coils to faithfully reproduce a desired pulse sequence, someembodiments of which are discussed in further detail below.

In some embodiments, the power component 1914 may be a “current mode”power component that drives a desired current through coil 2002. Thedesired current may be produced by power component 1914 in response to acurrent command from controller 1906. In this respect, the powercomponent 1914 may operate as a current source that is controlled by thecurrent command (which may be provided by the controller as a voltagelevel indicating the current to be provided to coil 2002). Controller1906 may change the current command such that power component 1914produces current values that change in accordance with a selected pulsesequence. For example, controller 1906 may command the power componentto drive one or more gradient coils in accordance with a pulse sequencecomprising a plurality of gradient pulses. For each gradient pulse, thepower component may need to ramp up the current provided to acorresponding gradient coil at the rising edge of the gradient pulse andramp down the current provided to the gradient coil at a falling edge ofthe gradient pulse. Example operation of a power component configured todrive the gradient coil to provide a plurality of such gradient pulsesis described in further detail below.

FIG. 21A shows an example of a gradient coil current waveform, accordingto some embodiments. In this example, the gradient coil current rapidlyramps up at the rising edge of the gradient pulse from 0 A to +20 Awithin a time interval of 0.2 ms, remains at +20 A for a period of time,then rapidly ramps down at the falling edge of the gradient pulse to −20A, and remains at −20 A for a period of time. It should be appreciatedthat the exemplary current to produce a gradient pulse is provided byway of illustration and that different pulse sequences may comprisegradient pulses having different current and/or voltage requirements.Controller 1906 and power component 1914 can be configured to drive oneor more gradient coils according to any suitable pulse sequence.

FIG. 21B shows waveforms for the current command, the gradient coilcurrent and the gradient coil voltage before, during and after therising edge of the gradient coil current shown in FIG. 21A. The gradientcoil current is the current through the gradient coil. The gradient coilvoltage is the voltage across the gradient coil. The current command isa signal representing an amount of current to be driven through thegradient coil by power component 1914. In response to a current commandat a time of 0 ms, the current through the gradient coil begins to risetoward the commanded current of +20 A. Since the gradient coil is aninductive load, a relatively large voltage needs to be provided to thegradient coil to rapidly increase the current through the gradient coil.Providing a rapid increase in current through the gradient coil isdesirable in MRI applications, as providing fast transitions betweengradient coil current values can decrease acquisition times and may beneeded to implement certain pulse sequences. As should be appreciatedfrom the exemplary voltages and currents shown in FIGS. 21A and 21B, thepower component 1914 may have the capability of driving the gradientcoil with relatively high power.

As an example, a gradient coil may have an inductance of 200 μH and aresistance of 100 mΩ Since the rate of change of the current through thegradient coil is proportional to its inductance, a voltage of 100V needsto be provided to the gradient coil to increase its current at a rate of100 A/ms. However, once the gradient coil current levels off at 20 A,the voltage requirement drops substantially. At this point, since thecurrent is no longer changing, the voltage needed depends upon theresistance of the gradient coil. Since the resistance of the gradientcoil is 100 mΩ, the voltage needed to be provided to the gradient coilto maintain the current steady at 20 A is 2V, which is significantlylower than the voltage (100V) needed during the transition betweencurrent values. However, these values of current, voltage, inductanceand resistance are provided merely by way of example, as any suitablegradient coil designs may be used, which may have different values ofinductance and/or resistance. Further, other suitable values ofcurrents, voltages, transition timings, etc. values may be used and/orneeded to implement a given pulse sequence.

Since the resistance of the gradient coil may be relatively low (e.g.,less than 500 mΩ), in some embodiments the power component 1914 has arelatively low output impedance in order to efficiently supply thecommanded current. For example, according to some embodiments, the powercomponent 1914 comprises a linear amplifier configured to power one ormore gradient coils (e.g., to provide current to the one or moregradient coils in accordance with a desired pulse sequence). Toimplement a linear amplifier having a low output impedance, transistorsof suitable size may be used having low equivalent series resistanceand/or a number of transistors may be connected in parallel to produce alow resistance collectively. Interconnects may be designed to have arelatively low resistance. The output impedance of the linear amplifiermay, for example, be less than twice the impedance of the gradient coil,in some embodiments. In some embodiments, the voltage drop across thetransistors of the linear amplifier may be relatively low in operation,such as less than 5V, less than 2V, or less than 1V (and greater than0V). Using an amplifier with a relatively low output impedance may beparticularly helpful for driving current through a gradient coil, whichmay have substantial DC current. A low output impedance can improveefficiency and limit heating of the amplifier. Details of exemplarylinear amplifier implementations are discussed in further detail below.

FIG. 22A shows an example of a power component 1914 having a currentfeedback loop and a voltage feedback loop, according to someembodiments. Power component 1914 is configured to provide the currentneeded to drive one or more gradient coils in accordance with a desiredpulse sequence. As such, power component 1914 is designed to be a lownoise current source that can be precisely controlled to provide thecommanded current waveform needed to drive the one or more gradientcoils to faithfully produce the desired gradient magnetic fields. Powercomponent 1914 includes a comparator 2101 that receives a currentcommand from controller 1906 at its non-inverting input terminal and acurrent feedback signal FB from a current sensor 2201 at its invertinginput terminal. The current command may be a voltage value representingthe commanded current. The current feedback signal FB may be a voltagevalue representing the measured current. In some embodiments, ahigh-quality current sensor may be used to provide an accurate feedbacksignal FB, which can improve the accuracy of the gradient coil currentpulses.

The comparator 2101 produces an error signal E (e.g., a voltage)representing the difference between the current command and the currentfeedback signal FB. Amplifier circuit 2102 amplifies the error signal toproduce an amplified error signal that is provided to the output stage2103. The output stage 2103 drives coil 2002 based upon the amplifiederror signal. The current through the coil 2002 is measured by currentsensor 2201, and a feedback signal FB is fed back to the comparator2101, as discussed above. The current feedback loop thereby causes thecurrent through the coil 2002 to be equal to the current commanded bythe controller 1906. In this respect, the power component 1914 mayoperate as a voltage-controlled current source. According to someembodiments, a high accuracy, high precision current sensor 2201 is usedto ensure that the current output provided to the gradient coilaccurately tracks the current commanded by the controller 1906. As aresult, the current provided to power the gradient coil can be held asclose to the commanded current as feasible. The power component 1914also has a voltage feedback loop that provides the output voltage of theoutput stage 2103 to the input of the voltage amplifier circuit 2102.

As illustrated in FIG. 22B, the voltage amplifier circuit 2102 mayinclude an operational amplifier OA that receives the error signal E atits non-inverting input and the voltage feedback signal V_FB at itsinverting input. The voltage feedback signal may be provided to theinverting input of the operational amplifier through a resistive voltagedivider (e.g., including resistors R1 and R2), which causes theoperational amplifier to amplify the input voltage based on the ratio ofresistance values in the voltage divider. Any suitable voltage gain maybe used for the voltage amplifier, such a gain of 5-15, by way ofexample. In some embodiments, the voltage gain of the output stage maybe one (unity).

As illustrated in FIG. 22A, in some embodiments, the controller 1906 mayprovide a command to the output stage 2103. The controller 1906 maycommand the output stage 2103 to produce a power supply voltage suitablefor supplying current needed to perform a corresponding portion of apulse sequence. As an example, the command may cause a power converterof the output stage to begin ramping up the magnitude of a power supplyvoltage in advance of a gradient coil current pulse. Such a command isdiscussed in further detail below with reference to FIG. 33D.

In some embodiments, the output stage 2103 is configured to beselectively powered by a plurality of power supply terminals atdifferent voltages. The power supply terminal selected to power theoutput stage 2103 may be chosen depending on the output voltage producedby the voltage amplifier. For example, when the power component iscommanded to produce a relatively high (positive) output voltage thepower component may be powered from a relatively high (positive) voltagesupply terminal, and when the power component is commanded to produce arelatively low (positive) output voltage, the power component is poweredfrom a relatively low (positive) voltage supply terminal. Accordingly,the efficiency of the power component can be improved by reducing thevoltage drop across its transistor(s) when relatively low output voltageis produced. It should be appreciated that any number of supplyterminals and voltage levels may be used, as the aspects are not limitedin this respect. For example, high, mid and low voltage supply terminals(both positive and negative) may be used, or an even greater number assuitable for a particular design and/or implementation.

FIG. 23A shows an example of an output stage 2103A having an outputVout, Iout suitable for powering one or more gradient coils of amagnetic resonance imaging system. To improve the power efficiency inpowering one or more gradient coils, output stage 2103A can be poweredby different supply terminals depending on the output voltage Vout. Forexample, output stage 2103A can be powered by a plurality of supplyterminals of a first polarity (e.g., a plurality of different positivevoltages) and/or a plurality of supply terminals of a second polarity(e.g., a plurality of different negative voltages). To facilitate lownoise operation, according to some embodiments, output stage 2103A mayinclude a linear amplifier 2104. According to some embodiment, each ofthe different supply terminals provides a different fixed supplyvoltage. According to some embodiments, one or more of the differentsupply terminals produce a variable supply voltage, as discussed infurther detail below.

In operation, if a positive output voltage is produced at Vout,switching circuitry S1 connects the high side power input of linearamplifier 2104 to either the high voltage terminal +Vhigh or the lowvoltage terminal +Vlow depending on the magnitude of the output voltage.If a relatively high output voltage is to be produced (e.g., if theoutput voltage to be produced exceeds a particular threshold), theswitching circuitry S1 connects the high side power input of linearamplifier 2104 to the high voltage terminal +Vhigh. If a relatively lowoutput voltage is to be produced (e.g., if the output voltage to beproduced remains below the particular threshold), the switchingcircuitry S1 connects the high side power input of linear amplifier 2104to the low voltage terminal +Vlow. Similarly, if a negative outputvoltage is produced, switching circuitry S2 connects the low side powerinput of linear amplifier 2104 to either the high voltage terminal−Vhigh or the low voltage terminal −Vlow depending on the magnitude ofthe output voltage, as discussed above. Any suitable switching circuitryS1 and S2 may be used. Such switching circuitry may include a diode thatis passively switched and/or a transistor that is actively switched.

In some embodiments, the high-voltage or low-voltage terminals may bedirectly connected to the linear amplifier 2104, without an interveningswitch S1 or S2. For example, as shown by the exemplary output stage2103A′ illustrated in FIG. 23B, the high voltage terminals +Vhigh and−Vhigh may be directly connected to the linear amplifier 2104, and thelow voltage terminals +Vlow and −Vlow may be connected to the linearamplifier 2104 through respective switches S1 and S2. The linearamplifier 2104 may be designed such that it is powered by a low voltagesupply terminal unless its voltage is insufficient to supply the outputcurrent, in which case the linear amplifier 2104 is powered by the highvoltage supply terminal. It should be appreciated that the use of+−Vhigh and +−Vlow is merely exemplary and any number of voltages levelsmay be used to provide a desired output voltage. For example, one ormore intervening voltage levels between +−Vhigh and +−Vlow,respectively, may be used to produce the desired voltage levels.

FIG. 24 shows an example of an output stage 2103A having a plurality ofdrive circuits 2401-2404. Drive circuits 2401-2404 drive a linearamplifier 2104 that includes a plurality of transistor circuits2405-2408, each including one or more transistors. The linear amplifier2104 can be connected to the high voltage or low voltage supplyterminals depending on the output voltage to be produced.

When a low positive output voltage is to be produced, transistor(s) 2406are connected to the low voltage terminal +Vlow via switch circuitry S3.Transistor(s) 2405 are turned off by drive circuit 2401 to isolate thetransistors 2406 from the high voltage terminal +Vhigh. Drive circuit2402 drives transistor(s) 2406 as a linear amplifying element, based onthe input, to produce an amplified output using the low voltage terminal+Vlow as a source of current.

To provide a high positive output voltage, drive circuit 2401 turns ontransistor(s) 2405 to connect the high voltage terminal +Vhigh to thetransistors 2406. Switch circuitry S3 may be turned off to isolatetransistor(s) 2406 from the low voltage terminal +Vlow. Drive circuit2402 may drive transistor(s) 2406 fully on, such that transistor(s) 2405are connected to the output of output stage 2103A. Drive circuit 2401drives transistor(s) 2405 as a linear amplifying element, based on theinput, to produce an amplified output using the high voltage terminal+Vhigh.

Accordingly, the low voltage terminal +Vlow can be used to provide a lowoutput voltage and the high voltage terminal +Vhigh can be used toprovide a high output voltage. A negative output voltage may be providedsimilarly by drive circuits 2403 and 2404, transistor(s) 2407 and 2408,and switch circuitry S4. When a negative output voltage is produced,drive circuits 2401 and 2402 may turn off transistor(s) 2405 and 2406.Similarly, when a positive output voltage is produced, drive circuits2403 and 2404 may turn off transistor(s) 2407 and 2408.

Transistor(s) 2406 may operate as a linear amplifying element of linearamplifier 2104 for low output voltages and transistor(s) 2405 mayoperate as a linear amplifying element for high output voltages. In someembodiments, transistor(s) 2406 and 2405 may be biased such that for atransition region between low positive output voltages and high positiveoutput voltages, transistor(s) 2405 and 2406 both act as linearamplifying elements of linear amplifier 2104, i.e., they are neitherfully-on nor fully-off. Operating both transistors 2405 and 2406 aslinear elements during such transitions may facilitate linear amplifier2104 having a smooth and continuous transfer function. Transistors 2407and 2408 may operate similarly to transistors 2405 and 2406 to produce arange of negative output voltages.

In some embodiments, switch circuitry S3 and S4 may be realized bydiodes that automatically switch on an off depending on whether the highvoltage terminal is being utilized. For example, if switch circuitry S3includes a diode, the anode may be connected to the terminal +Vlow andthe cathode to transistor(s) 2406, such that current can only flow outof terminal +Vlow into the output stage 2103A. However, the techniquesdescribed herein are not limited in this respect, as switch circuitry S3and S4 may be realized using controlled switches, such as transistors,or any other suitable switching circuitry.

In some embodiments, the circuit of FIG. 24 may be used to drive agradient coil using a pulse sequence as shown in FIG. 21. When theoutput current is constant, the output voltage (e.g., 2V) may beproduced by sourcing current from the low voltage terminal +Vlow. Duringa transition when the current is changed rapidly, a high output voltage(e.g., 100V) may be produced by sourcing current from the high voltageterminal +Vhigh. Thus, the high voltage terminal may be used duringtransitions in the output current to provide high output voltages, andthe low voltage terminal may be used to provide low output voltages forhigh efficiency.

According to some embodiments, for example, according to some pulsesequences, the high voltage terminal(s) may only need to be used for arelatively short period of time, so that transistor(s) 2405 (and 2408)may be conducting for only a relatively small duty cycle. Thus, in someembodiments, transistor(s) 2405 (and 2408) may be reduced in size,and/or the number of transistors connected in parallel may be reduced,with respect to transistors 2406 (or 2407), as transistor(s) 2405 (and2408) will have time to dissipate heat between transitions in thegradient coil current.

In some embodiments, drive circuits 2401 and 2404 may be designed toprovide time-limited output signals. Providing time-limited outputsignals may ensure that transistor(s) 2405 and/or 2408 are turned ononly temporarily and not turned on to drive a steady state current. Sucha technique may be advantageous if transistor(s) 2405 or 2408 aredesigned to conduct for only relatively short periods of time, as it canprevent excessive power dissipation by transistor(s) 2405 or 2408.

FIG. 25 shows a block diagram of drive circuitry 2401 and 2402,according to some embodiments. Drive circuitry 2401 includes a drivetransistor 2503A for driving transistor(s) 2405. Drive circuitry 2402includes a drive transistor 2503B for driving transistor(s) 2406.

Drive circuitry 2401 and 2402 may include one or more bias circuits 2501for producing a DC bias on the input voltage provided to the drivetransistors 2503A and 2503B. In some embodiments, the bias circuit(s)2501 may bias drive transistors 2503A and/or 2503B slightly below theirturn-on voltages. The inventors have recognized and appreciated thatbiasing the drive transistors slightly below their turn-on voltages canreduce or eliminate thermal runaway. Advantageously, such a biasingtechnique may not reduce the linearity of the output stage 2103A. If anoperational amplifier OA of voltage amplifier circuit 2102 has asufficiently high speed, it can respond fast enough to accuratelycontrol the output voltage of the output stage despite biasing the drivetransistors slightly below their turn-on voltages.

In some embodiments, drive circuitry 2401 may include a timing circuitthat causes drive circuit 2401 to produce a time-limited output. Anysuitable timing circuit may be used. In the example of FIG. 25, a timingcircuit 2502 is connected to the input of output stage 2103A via biascircuit 2501, and limits the amount of time that an input can beprovided to the drive transistor 2503A.

In some embodiments, the timing circuit 2502 may be an RC circuit thathas an output voltage that decays over time, and turns off drivetransistor 2503A when the output of the timing circuit 2502 falls belowthe turn on voltage of the drive transistor 2503A. The time thattransistor(s) 2405 are turned on is limited based on the RC timeconstant of the RC circuit. However, the techniques described herein arenot limited to implementing the timing circuit using an RC circuit, asany suitable timing circuitry may be used, including analog and/ordigital circuitry. In some embodiments, drive circuits 2403 and 2404 maybe implemented similarly to drive circuits 2402 and 2401, respectively,for negative input and output voltages.

FIG. 26 shows an example implementation of the drive circuits of FIG.25, in accordance with some embodiments of the technology describedherein. As shown in FIG. 26, in some embodiments, the bias circuit 2501may be realized by a Zener diode in series with a resistor R2, connectedbetween the high voltage terminal +Vhigh and a lower voltage DC terminal(e.g., −Vhigh) below the voltage of +Vhigh. In some embodiments, thebias circuit 2501 may include additional circuitry between the highvoltage terminal +Vhigh and the lower voltage DC terminal to provide aDC path for current to flow between them and establish a suitable biasvoltage. In some embodiments, the bias circuit 2501 may include anotherZener diode and resistor in series with the Zener diode and resistorshown in FIG. 26, for providing bias voltage(s) to the low-side drivecircuits 2403 and 2404. However, this is merely by way of example, asany suitable bias circuit may be used. FIG. 26 also illustrates anexample of a timing circuit 2502 realized as an RC circuit having acapacitor C1 and a resistor R1. Again, this is merely one example of atiming circuit, as other configurations of timing circuits may be used.Drive transistors 2503A and 2503B are shown as being realized by bipolarjunction transistors. However, the techniques described herein are notlimited in this respect, as the drive transistors may be realized by anytype of transistors. Transistor circuits 2405 and 2406 are shown asMOSFETs, in this example. However, the transistor circuits 2405 and 2406may be realized by any type of transistors. In some embodiments,transistor circuits 2405 and/or 2406 may have a plurality of transistorsconnected in parallel. As discussed above, switch circuitry S3 may berealized as a diode, as shown in FIG. 26. However, as discussed above,the techniques described herein are not limited in this respect, as insome embodiments switch circuitry S3 may be realized by a transistor.

FIG. 27 shows another example of a technique for implementing a timingcircuit. The inventors have recognized and appreciated that if switch S3is realized by a diode, the voltage across the diode can be used as atrigger for a timing circuit to limit the amount of time thattransistor(s) 2405 are turned on. When a low output voltage is producedby linear amplifier 2104, the diode is forward biased and conducting.When the linear amplifier 2104 produces a high output voltage,transistor(s) 2405 turn on and the diode switches from being forwardbiased to being reverse biased. The reverse bias voltage can be sensedby timing circuit 2702 as an indication that transistor(s) 2405 arebeing turned on. In the example of FIG. 27, the voltage across the diodeis provided as an input to timing circuit 2702, which produces aninhibit signal to inhibit the operation of drive circuit 2401 after aperiod of time, thereby limiting the amount of time that transistor(s)2405 are turned on. Timing circuit 2704 may operate similarly in orderto inhibit the operation of drive circuit 2404 after transistor(s) 2408have been conducting for a period of time.

FIG. 28 shows an example of timing circuits 2702 and 2704 realized by anRC circuit and a bipolar transistor. In timing circuit 2702, forexample, once the diode is reverse biased after a period of time theoutput of the RC circuit rises to a level where the bipolar transistorturns on. When the bipolar transistor turns on, the input of the drivecircuit 2401 is pulled down to +Vlow, which turns off the drive circuit2401 and transistor(s) 2405.

Although FIGS. 24, 27 and 28 show a “double-ended” linear amplifier 2104that may produce a positive output voltage or a negative output voltage,the techniques described herein are not limited in this respect, as insome embodiments a single-ended linear amplifier may be used. FIG. 29shows an example of an output stage 2103B including a single-endedlinear amplifier 2105 that produces only positive output voltages. FIG.29 schematically illustrates that the single-ended linear amplifier 2105may be connected to a high positive voltage terminal +Vhigh or a lowpositive voltage terminal +Vlow by switch S1, depending on the outputvoltage to be produced. The output stage 2103B may be implemented usingthe drive circuits 2401, 2402, transistor(s) 2405 and 2406, andassociated switch circuit S3 discussed above, in some embodiments.

The output stage 2103B may provide a positive output voltage or anegative output voltage to a load using a polarity-switching circuit2904. In the example of FIG. 29, polarity-switching circuit 2904 isrealized using an H-bridge including switches S5-S8. A positive voltagemay be provided to the load by turning on switches S5 and S8 and turningoff switches S6 and S7. A negative voltage may be provided to the loadby turning on switches S6 and S7 and turning off switches S5 and S8. Insome embodiments, the control circuit (not shown) may control switchesS5-S8 to produce an output voltage of a suitable polarity. The polaritymay be determined by examining the polarity of the current command, theerror signal E, or any other suitable signal.

As discussed above, conventional switching converters can introduce asignificant amount of switching noise into the system because theyswitch at frequencies in the range of tens to hundreds of kHz. Suchswitching noise can interfere with imaging because it is in the samefrequency range as MR signals desired to be detected. The inventors haverecognized that a power converter having a switching frequency above theLarmor frequency of interest does not interfere with imaging to asignificant degree. Accordingly, in some embodiments, power component1914 may include a switching power converter 3002 that is designed toswitch at a relatively high switching frequency, above the Larmorfrequency of interest, as shown in FIG. 30. In some embodiments, theswitching frequency may be higher than 1 MHz, higher than 10 MHz, higher30 MHz or higher than 300 MHz.

As discussed above, the inventors have appreciated that providingvariable voltage supply terminals facilitates efficient powering of oneor more gradient coils of a magnetic resonance imaging system (e.g., alow-field MRI system). In some embodiments, the output stage may bepowered by one or more or more variable voltage supply terminals thatare controlled to produce supply voltages close to the desired outputvoltage. Providing such a variable voltage supply terminal can improvethe efficiency of the output stage by limiting the voltage drop acrossthe linear amplifier.

FIG. 31 shows an embodiment of an output stage 2103C that may be poweredby a variable voltage positive supply terminal and a variable voltagenegative supply terminal. The voltages of the supply terminals can bevaried depending on the output voltage to reduce the voltage drop acrossthe transistor(s) of the linear amplifier 2106, thus facilitatingefficient powering of gradient coil(s) to produce magnetic fieldsaccording to a desired pulse sequence. In some embodiments, the voltageof the positive voltage terminal and/or the negative voltage terminalmay be provided by power converters 3104 and/or 3106, respectively. Thevariable output voltages of the power converters 3104 and/or 3106 may becontrolled by a controller 3108 based on the desired output voltage ofoutput stage 2103C to maintain the voltages of the positive voltageterminal and/or the negative voltage terminal slightly above (or below,respectively) the output voltage of the output stage, thereby reducingthe voltage drop across the transistor(s) of the linear amplifier.

According to some embodiments, controller 3108 controls the variableoutput voltages of the power converters 3104 and/or 3106 based on theoutput voltage of linear amplifier 2106. However, the variable outputvoltages may be controlled in other ways and/or in differentrelationship to the desired output voltage of output stage 2103C. Forexample, the variable output voltages may be controlled based on thecommand (e.g., current command) provided to linear amplifier 2106. Asdiscussed in the foregoing, a controller may be configured to commandthe linear amplifier to produce output sufficient to drive one or moregradient coils of a magnetic resonance imaging system in accordance witha desired pulse sequence. As such, controller 3108 may be configured tocontrol the variable output voltages of the power converters 3104 and/or3106 so that the output voltages provided to the linear amplifier aresufficient, without being too excessive and therefore inefficient, toallow the linear amplifier to produce output to power the one or moregradient coils in accordance with the desired pulse sequence. Control ofthe power converters 3104 and 3106 may be performed in any suitable way,such as by controlling their duty ratio, their frequency, or any othercontrol parameter that can control the output voltage of the powerconverters. In some embodiments, power converters 3104 and 3106 of FIG.31 may be switching power converters designed to switch at a relativelyhigh switching frequency, above the Larmor frequency of interest, asdiscussed above. However, any suitable power converter may be used, asthe aspects are not limited in this respect.

In some embodiments, both high and low voltage supply terminals (e.g.,+Vhigh and +Vlow) may power the linear amplifier, as illustrated inFIGS. 23, 24, and 29, and the voltage of the low voltage supplyterminal, the high voltage supply terminal, both, or any supply terminalprovided may be variable. FIG. 32A shows an embodiment of an outputstage 2103D similar to FIG. 23A with variable low voltage supplyterminals. Rather than having low voltage terminals +Vlow and −Vlow atfixed voltages, FIG. 32A shows that +Vlow and −Vlow can have variablevoltages. In some embodiments, the variable voltages of +Vlow and −Vlowmay be provided by power converters 3203 and 3204, respectively. In someembodiments, power converters 3203 and 3204 may be switching powerconverters designed to switch at a relatively high switching frequency,above the Larmor frequency of interest, as discussed above. When arelatively low output voltage is to be produced (e.g., in the steadystate), current is sourced from the low voltage supply terminals +Vlowor −Vlow. The output voltages +Vlow or −Vlow of the power converters3203 or 3204 may be controlled by controller 3108 based on the desiredoutput voltage Vout of linear amplifier 2104 to maintain the voltages ofthe low voltage supply terminals +Vlow or −Vlow slightly above (orbelow, respectively) the output voltage of the output stage, therebyreducing the voltage drop across the transistor(s) of the linearamplifier in the steady state and reducing power dissipation. When arelatively high output voltage is to be produced, current may be sourcedfrom the high voltage terminals +Vhigh or +Vlow, which may have fixedvoltages.

+Vhigh may be a separate terminal from the power supply terminalVhigh_Supply that supplies power to power converter 3203, as illustratedin FIG. 32A, or may be the same terminal as Vhigh_Supply, as illustratedin FIG. 32B. In FIG. 32B, an example is shown of an output stage 2103Ein which +Vhigh is provided from the power supply terminal Vhigh_Supplyand −Vhigh is provided from the power supply terminal Vlow_Supply thatprovides power to power converter 3204. Providing +Vhigh and/or −Vhighfrom the existing power supply terminals can avoid the need to produceadditional power supply voltages, which can simplify the design andimplementation of the output stage.

FIG. 33A shows an example of a gradient coil current waveform, accordingto some embodiments. The gradient coil current is initially zero, thenrapidly ramps up to 10 A in 0.1 ms. The current remains at 10 A for aperiod of time, then drops back to 0 A. The current remains at 0 A for aperiod of time before rapidly ramping up to 20 A in 0.2 ms. The currentremains at 20 A for a period of time, then drops back to 0 A. It shouldbe appreciated that the amp values and time intervals are merelyexemplary for the purposes of illustration, and any suitable values maybe used.

FIG. 33B shows the rising transition of the gradient coil current from 0A to 10 A, the voltage 3302 needed for driving the gradient coil, thevoltage of the high voltage supply terminal +Vhigh and the low voltagesupply terminal +Vlow. During the transition, current is sourced fromthe high voltage supply terminal +Vhigh in order to provide a highvoltage to the gradient coil to quickly ramp up its current. As thetransition occurs, the power converter 3203 begins to ramp up thevoltage of the low voltage supply terminal +Vlow from ˜0V to a voltageslightly higher than the output voltage necessary to drive the gradientcoil with a steady state current of 10 A. Once the steady state currentof 10 A is reached, current is sourced from the low voltage supplyterminal +Vlow in order to provide high efficiency in the steady state.

FIG. 33C shows the rising transition of the gradient coil current from 0A to 20 A, the gradient coil voltage, and the voltage of the highvoltage supply terminal +Vhigh and the low voltage supply terminal+Vlow. During the transition to 20 A, as with the transition to 10 A,current is sourced from the high voltage supply terminal +Vhigh in orderto provide a high voltage to the gradient coil to quickly ramp up itscurrent. As the transition occurs, the power converter 3203 begins toramp up the voltage of the low voltage supply terminal +Vlow from ˜0V toa voltage slightly higher than the output voltage necessary to drive thegradient coil with a steady state current of 20 A. Once the steady statecurrent of 20 A is reached, current is sourced from the low voltagesupply terminal +Vlow.

Since the voltage of the low voltage supply terminal +Vlow can bevaried, it can be set slightly above the output voltage needed fordifferent steady state current levels. This can improve the efficiencyover the case of using a low voltage supply terminal +Vlow having afixed voltage, as a fixed voltage would need to be designed to handlethe maximum steady state current, which may be a higher voltage thannecessary for driving lower steady state currents, which can decreaseefficiency. As an example, if the +Vlow is set high enough to supply a20 A steady state gradient coil current, such a voltage is higher thannecessary to supply a 10 A steady state gradient coil current, whichresults in increased voltage drop across the linear amplifiertransistor(s) when supplying a 10 A steady state gradient coil current,and higher power dissipation occurs than is necessary. A variablevoltage can be set at or near the minimum voltage necessary to supplythe commanded steady state gradient coil current, which improvesefficiency.

FIG. 33D shows a current command, gradient coil current, the voltage3302 of the gradient coil needed to supply the current, and severaldifferent transition waveforms of the voltage +Vlow. Transition waveform3304 shows an idealized transition in which the voltage of +Vlow startsramping up in response to the rising edge of the gradient coil currentcommand, and reaches the steady state value of +Vlow at the same timethat the steady state gradient coil current (and voltage values) arereached. However, the inventors have recognized and appreciated thatthere may be factors preventing the voltage +Vlow from reaching asufficient voltage level in time for the terminal +Vlow to supply thesteady state current. Transition waveform 3306 shows a more realistictransition of +Vlow, which has a period of latency (delay) in respondingto the gradient coil current command. As shown in FIG. 33D, thetransition waveform 3306 starts ramping up only after a period of timefollowing the rising edge of the current command. The slope of thetransition waveform 3306 may be limited, as the power converter 3203 mayhave an output filter (e.g., a capacitor) that limits the speed withwhich power converter 3203 can change the voltage of +Vlow. As a result,the transition waveform 3306 may not reach a sufficient voltage level bythe time the steady state gradient coil current and voltage are reached,which may result in the low voltage supply terminal +Vlow being unableto supply the steady state current, at least temporarily.

To address this, in some embodiments, the power converter 3203 (or 3204)may begin ramping up the magnitude of the voltage of +Vlow (or −Vlow)before the rising edge of the gradient coil current command. FIG. 33Dshows a transition waveform 3308 for +Vlow that starts ramping up beforethe rising edge of the gradient coil current command. To begin thetransition prior to the rising edge of the gradient coil currentcommand, controller 3108 may receive information from controller 1906regarding an upcoming gradient coil current pulse, and begin ramping upthe magnitude of the voltage of +Vlow (or −Vlow) in anticipation of thecurrent pulse. This information may be provided from controller 1906 tocontroller 3108 in any suitable way. As an example, the controller 1906may analyze the currently selected gradient coil pulse sequence,determine a power supply voltage level suitable to supply the steadystate gradient coil current for the next current pulse, and send avoltage command to controller 3108 in advance of an anticipated currentcommand. The power converter 3203 (or 3204) may then respond to thereceived voltage command and begin ramping +Vlow (or −Vlow) to thecommanded voltage value. As another example of providing the informationto the controller 3108, the controller 1906 may send the currentlyselected pulse sequence or a portion of the pulse sequence to controller3108. Controller 3108 may then analyze the pulse sequence and sendcommands to the power converter 3203 (or 3204) to start ramping thevoltage +Vlow (or −Vlow) in advance of a gradient coil current pulse. Inthe example of FIG. 33D, the power converter 3203 starts ramping up thevoltage of +Vlow in response to a voltage command provided by controller1906 to controller 3108 in advance of the rising edge of the currentcommand As a result, the transition waveform 3308 reaches the level of+Vlow sufficient to supply the steady state current by the time thesteady state current level is reached.

FIG. 34A shows an embodiment of an output stage 2103F with asingle-ended linear amplifier similar to that of FIG. 29, with avariable low voltage supply terminal +Vlow. As with the embodiment ofFIG. 32A, the power converter 3203 supplies a variable voltage to thelow voltage supply terminal +Vlow that can be set slightly above thevoltage needed to supply the commanded steady state gradient coilcurrent.

As discussed above in connection with FIG. 32A and FIG. 32B, the highvoltage supply terminal +Vhigh may be a separate terminal from the powersupply terminal Vhigh_Supply, as illustrated in FIG. 34A, or may be thesame terminal as Vhigh_Supply, as illustrated in FIG. 34B. In FIG. 34B,an example of an output stage 2103G is shown in which +Vhigh is providedfrom the power supply terminal Vhigh_Supply. Providing the voltage+Vhigh from the existing power supply terminal Vhigh_Supply can avoidthe need to produce additional power supply voltages, which can simplifythe design and implementation of the output stage.

In some embodiments, both the low voltage supply terminal(s) and thehigh voltage supply terminal(s) may have variable voltages. For example,the embodiments of FIG. 32 or 29 may be modified such that the highvoltage supply terminals +Vhigh and/or −Vhigh are variable voltagesproduced by power converters. Such power converters may be similar topower converters 3203 and 3204, and may be controlled by the controller3108, as well. Such an embodiment can be used for any suitable type ofimaging, and may be particularly useful for diffusion weighted imaging,for example, where relatively large currents may be needed (e.g., 40A,50A, 70A, 90A or more, or any values there between).

In some embodiments, one or more additional power supply terminals maypower the linear amplifier. For example, a third power supply terminalmay be provided that has a voltage higher than the high voltage supplyterminal +Vhigh (e.g., at least 5 times higher or at least 10 timeshigher, and even as high as 20 or 30 times higher or more, or in anyrange between such values). Adding a third supply terminal may helpimprove efficiency in the case where a wide range of voltages need to beproduced. Any number of power supply terminals may be provided, as thetechniques described herein are not limited in this respect.

Accordingly, using techniques described herein for a low power, lownoise amplifier, gradient amplifiers may be configured to operate wellwithin the power budget available using mains electricity (e.g., powerdelivered from standard wall outlets. According to some embodiments thatutilize a linear amplifier design, the power electronics for poweringthe gradient coils of an MRI system consume between 100-200 W fortypical pulse sequences (e.g., bSSFP, FLAIR, etc.) and between 200 W-750W for more demanding pulse sequences such as DWI. According to someembodiments using switched power converters, the power electronics forpowering the gradient coils of an MRI system consume between 50-100 W orless for typical pulse sequences (e.g., bSSFP, FLAIR, etc.) and between100 W-300 W or less for significantly demanding pulse sequences such asDWI. In addition to low power operation that facilitates operation usingstandard wall power, the gradient power amplifiers described herein arealso configured to be relatively compact in size so that they can behoused within an enclosure (e.g., within base 1950 of the portable MRIsystems described in FIGS. 19A and 19B) along with the other electroniccomponents to facilitate portable MRI. According to some embodiments,the gradient amplifiers are designed to be connected to a backplane(e.g., a printed circuit board backplane) that connects the gradientamplifiers to the power source (e.g., wall power) and to the gradientcoils of the system, as discussed in further detail below in connectionwith FIGS. 36 and 37A-D.

The inventors have further developed low power and efficient amplifiersto operate the RF coils of the RF transmit/receive system (e.g., RFpower amplifiers to drive one or more transmit/receive coils configuredto produce B₁ magnetic field pulses to produce an MR response) tofacilitate operation of a portable MRI system. According to someembodiments, RF power amplifiers (RFPAs) are configured to operate usingmains electricity (e.g., sharing a portion of available main electricitywith the other components of the system), such as the power suppliedfrom a single-phase standard wall outlet and/or from a single-phaselarge appliance outlet. In embodiments of a portable MRI systemoperating with power supplied from single phase wall power, the RFPAsmust share the limited available power with other components (e.g., theexemplary GPAs discussed above, console, on-board computer, coolingfans, etc.) and therefore need to be designed to efficiently make use ofthe limited power available. The inventors have developed techniques forefficient RFPAs suitable for use in portable MRI powered by mainselectricity. According to some embodiments, the maximum input power tothe RFPA(s) is approximately 160 W, thereby limiting the average powerconsumption of the RFPA(s) to a maximum of 160 W. However, thetechniques described herein significantly reduce the average powerconsumption of the RFPA(s), including in circumstances when a givenpulse sequence requires higher levels of instantaneous power (e.g., 400W for DWI pulse sequences), as discussed in further detail below.

FIG. 35 is a block diagram of an exemplary low power RFPA, in accordancewith some embodiments. RFPA 3500 comprises an input block 3510 thatreceives an RFIN signal 3502 corresponding to the desired RF signalwaveform to be amplified by power amplifier 3550 and provided as RFOUT3522 at power levels sufficient to operate an RF transmit coil toproduce B₁ magnetic field pulses according to a desired pulse sequence.Power amplifier 3550 may include any suitable type of amplifier, orcombination of amplifier stages, to amplify RFIN 3502 to suitablelevels. For example, power amplifier 3550 may comprise one or more classA type amplifiers configured to amplify RFIN 3502 to a desired powerlevel RFOUT 3522 (e.g., maximum 100 W, 400 W, etc. of instantaneouspower). Class A amplifiers provide excellent fidelity in amplifying aninput signal and therefore facilitate producing an RFOUT waveform 3522with very little distortion, ensuring that image quality is not degradedby distortion from the RFPA. However, other classes of amplifiers aremore power efficient, but generally increase distortion in the course ofamplifying an input signal. The inventors have recognized that due tothe relatively high Q factor of the RF coils, some additional distortionin the RFOUT 3522 may be tolerated with little, no or acceptable impacton image quality. According to some embodiments, power amplifier 3550comprises one or more class B, class AB or class BC type amplifiers,etc. Thus, power amplifier 3550 may be made more power efficient byselecting a more efficient class of amplifier provided the increaseddistortion can be tolerated or compensated for so that image quality isnot unsatisfactorily sacrificed. According to some embodiments, poweramplifier 3550 includes a plurality of amplification stages thatincrementally step up the signal to the desired level for RFOUT 3522.

RFPA 3500 comprises a power entry module 3572 that receives power atpower-in 3570, which may correspond to power at a number of differentpower levels. The power provided at power-in 3570 may be provided by theMRI system's power supplies that deliver DC power converted from AC wallpower, as discussed in further detail below in connection with FIG. 36.It should be appreciated that one or more DC-DC power supplies may alsobe provided to produce desired voltage levels from the DC power providedby the AC-DC power supply or supplies. For example, power-in 3570 mayinclude power lines at +100V, +40V, +23V and −15V provided by, forexample, a switched power supply that receives DC power from the AC-DCpower supply or supplies. According to some embodiments, a power supplyboard is included to provide the voltage levels needed for the RFPA. Thepower supply board may be implemented as a card that is connected to abackplane to receive DC power from the AC-DC power supply and convertthe power into desired voltage levels that are delivered to RFPA, whichitself may implemented as a board connected to the same backplane, asdiscussed in further detail below in connection with FIG. 36. Powerregulation module 3575 includes regulators that convert the powerreceived by the power entry module 3572 to the voltage levels needed bythe RFPA. For example, power regulation module 3575 may include powerregulator(s) that provide power lines at +/−5V, +13.8V, +15V and +3.3Vto be distributed to controller 3560, power amplifier 3550 and/or anyother components of the RFPA requiring power. It should be appreciatedthat the power distribution arrangement and the power levels needed willdepend on the requirements of the particular system, and the abovementioned values are merely exemplary.

In conventional MRI systems, RFPAs typically consume the maximum powerrequired to transmit the B₁ magnetic pulse sequences continuously. Inparticular, even when maximum power is not required for a particularpulse sequence and during intervals when no RF pulses are being produced(e.g., during a transmit quiet period when the MRI system is detectingemitted MR signals), conventional RFPAs still consume maximum power.Because conventional MRI systems are generally not power limited (e.g.,conventional MRI systems are powered by a dedicated three-phase powersources), the inefficient use of power consumption is generallyacceptable and tolerated. However, an RFPA consuming maximum power maybe unsuitable for low power MRI, for example, for a portable MRI systemoperating from the power supplied by mains electricity (e.g.,single-phase wall power). The inventors have developed techniques formore optimal operation of an RFPA from a power consumption perspective,facilitating operation of a portable MRI system using mains electricity.

In FIG. 35, a controller 3560 (e.g., a field programmable gate array(FPGA) is provided to control various aspects of the operation of RFPA3500 to reduce power consumption and/or more efficiently use availablepower, examples of which are described in further detail below.According to some embodiments, an RFPA is configured so that the maximumpower drawn by the power amplifier is selectable based on the powerrequirements of a given pulse sequence. In particular, different imageacquisition pulse sequences have different power requirements. Forexample, a diffusion weighted imaging (DWI) pulse sequence requiressignificantly more power than a balance steady-state free precession(bSSFP) pulse sequence. Conventionally, RFPAs would be set to providethe power amplification needed for the most demanding pulse sequences(e.g., to draw power according to the most power intensive pulsesequences), such as DWI pulse sequences for example. As a result, duringimage acquisition of less demanding pulse sequences (e.g., bSSFP),significant excess power is consumed by the RFPAs.

RFPA 3500 is configured so that the power amplification can be selectedbased upon the requirements of a given pulse sequence (e.g., the powerdissipation of the power amplifier can be selected based on the powerneeded to produce a given pulse sequence). In exemplary RFPA 3500illustrated in FIG. 35, a power select signal 3508 may be provided tocontroller 3560 to configure power amplifier 3550 to amplify RFIN 3502in accordance with the maximum power requirements of a given pulsesequence. For example, for a pulse sequence that requires 50 W, thepower select signal 3508 may instruct controller 3560 to bias poweramplifier 3550 to dissipate 50 W. Similarly, for a pulse sequence thatrequires 100 W, the power select signal 3508 may instruct controller3560 to bias power amplifier 3550 to dissipate 100 W, and for a pulsesequence that requires 400 W, the power select signal 3508 may instructcontroller 3560 to bias power amplifier 3550 to dissipate 400 W, etc. Inthis manner, power amplifier 3550 may be scaled to dissipate power inproportion to the maximum power needs of a given pulse sequence. Thus,because RFPA 3500 is not always consuming power according to the maximumpower requirements of the most demanding pulse sequence, significantpower reduction may be achieved. According to some embodiments, thepower select signal 3508 is set by the console based on the pulsesequence to be used to perform a given image acquisition protocol.

While the power select signal 3508 allows scaling of the power amplifierto the maximum power requirements of a given pulse sequence, excesspower will still be consumed during intervals where the pulse sequencedoes not require maximum power levels, thereby reducing the possibleefficiency of the RFPA. To address this drawback, the inventors havedeveloped techniques to dynamically scale power dissipation of the poweramplifier in accordance with the changing needs of a given pulsesequence. According to some embodiments, the power consumed by the RFPAis dynamically adjusted based on the needs of the signal beingamplified. For example, as illustrated in FIG. 35, controller 3560 mayreceive an envelope signal 3504 corresponding to the amplitude of theRFIN 3502 waveform to provide an indication of the instantaneous powerlevels needed to produce the desired RF pulse sequence. Based onenvelope 3504, controller 3560 may be configured to dynamically biasamplifier 3550 in correspondence to the changing envelope of the RFINwaveform (e.g., by changing the biasing points on the amplifiertransistors in correspondence with envelope signal 3504). As a result,the envelope or magnitude of RFIN 3502 may be tracked by controller 3560via envelope signal 3504 to dynamically bias the power amplifieraccordingly, thus limiting the power dissipation of power amplifier 3550to the contemporaneous power needs of the pulse sequence andsignificantly reducing excess power consumed by RFPA 3500. In thismanner, power amplifier 3550 can be scaled to draw power in accordancewith the instantaneous or substantially instantaneous power needs of thetransmitted pulse sequence.

As discussed above, a pulse sequence typically defines the timing ofboth RF and gradient magnetic field pulses as well as defining the timeperiods during which the receive coils are detecting MR pulses (e.g.,so-called transmit quiet periods). Thus, pulse sequences will haverepeated intervals of time when no RF magnetic field pulses are beingtransmitted. The inventors recognized that if the power amplifierremains on during these intervals (e.g., during transmit quiet periods),power will be consumed even though no RF magnetic field pulses are beingtransmitted. According to some embodiments, one or more power consumingcomponents of the RFPA are turned off during periods when no RF magneticfield pulses are being produced by the RF transmit coil(s) (e.g., duringtransmit quite periods such as during MR signal detection and/or duringsome portions of gradient pulse sequences generation) to prevent theRFPA from consuming power unnecessarily.

As an example, in exemplary RFPA 3500, controller 3560 receives anunblanking signal 3506 that indicates transmit quiet periods of thecurrent pulse sequence. In response to the unblanking signal 3506indicating a transmit quiet period, controller 3560 is configured toturn off power amplifier 3550 to the extent possible to conserve power(e.g., logic and bias circuits and any other circuitry that consumespower and can be turned off or disconnected may be shut down bycontroller 3560). When the unblanking signal 3506 changes state toindicate that an RF magnetic field pulse is to be produced, controller3560 turns on the power amplifier 3550 so that the RF magnetic fieldpulse can be produced and transmitted by the RF coil(s). Unblankingsignal 3506 may be provided by the console or main controller of the MRIsystem to indicate transmit quiet periods of the pulse sequence of animage acquisition procedure. In many pulse sequences, intervals when RFmagnetic field pulses are transmitted may be as little as 10% of thepulse sequence. As such, disabling the power amplifier during thesignificant transmit quiet periods may result in relatively significantpower savings.

It should be appreciated that one or a combination of the abovedescribed techniques may be used to reduce the power consumption of theRFPA to facilitate low power MRI, as the aspects are not limited in thisrespect. In particular, an RFPA need not include each of the powersaving techniques described above, but instead can employ one or more ofthese techniques. For example, a RFPA may include a mechanism thatallows selection of discrete power levels depending on the pulsesequence (e.g., via power select signal 3508), a mechanism to scale thepower of the power amplifier according to the instantaneous (orapproximately instantaneous) power needs of the RF pulses (e.g., bytracking the envelope 3504 of the RF pulse waveform) and/or a mechanismto disable the power amplifier during transmit quite periods (e.g., viaunblanking signal 3506). Using one or more of the techniques describedabove, the RFPA(s) may consume less than the 160 W input power even whenproducing demanding pulse sequences such as DWI that require intervalsof instantaneous power that exceed the input power (e.g., 400 W ofinstantaneous power). According to some embodiments, the RFPA(s) of aportable MRI system consume 65 W or less during image acquisition and,according to some embodiments, the RFPA(s) may consume 50 W or less(e.g., 25-30 W or less) during operation depending on the pulsesequences produced, thus conserving available wall power for the othercomponents of the system (e.g., GPAs, computer, console, fans, etc.).Other power savings techniques may be used in addition or in thealternative, as the aspects are not limited in this respect.

The above described components facilitate low power operation of an MRIsystem allowing for a portable MRI system that can be operated usingmains electricity (e.g., single-phase “wall power” delivered at standardand/or large appliance outlets). In addition to low power consumption,aspects of portability of an MRI system may be enhanced by a compactdesign where electronic components used to operate the MRI system arecontained on or within a standalone unit along with the magneticscomponents of the system. Incorporating the power conversion anddistribution system, the electronic components (e.g., power amplifiers,console, on-board computer, thermal management, etc.) and the magneticscomponents of the MRI system on or within a single self-containeddevice, facilitates portable MRI. As discussed above, conventional MRIsystems typically have a separate room for the power components, whichmust therefore deliver power to the magnetics components of the MRIsystem via cables connecting the power components to the MRI devicelocated in a specially shielded room. Not only is this arrangement fixedto a dedicated space, but the cabling required to connect the powercomponents to the magnetics components is responsible for significantpower losses. As discussed above in connection with FIGS. 19A and 19B,to facilitate portability, the inventors have developed a power systemthat is contained within a housing that supports or on which themagnetics components of the MRI system are located to provide astandalone, portable MRI system that can be brought to any locationhaving access to wall power, some examples of which are described infurther detail below.

FIG. 36 illustrates a block diagram of components of a portable MRIsystem 3600, including the magnetics components (e.g., B₀ magnet 3622,gradient coils 3628, RF coil(s) 3626 and, optionally, shim coils 3624),a power conversion and distribution system configured to receive powerfrom a mains electricity source (e.g., a single-phase wall outlet) andthe electronic components used to operate the magnetics components andcontrol the operation of the portable MRI system. In FIG. 36, portableMRI system 3600 includes an electronics enclosure 3602, a preamplifierenclosure 3604 and a fan board enclosure 3606. Electronics enclosure3602 may be positioned below and/or arranged to support the magneticscomponents of the MRI system (e.g., the B₀ magnet, gradient coils, RFcoils, shim coils, etc.) to provide a single, integrated standalone andportable MRI system (e.g., enclosure 3602 may form, in part, a base forthe portable MRI system 3600, similar to base 1950 located below themagnetics components in portable MRI system 1900 illustrated in FIGS.19A and 19B). Electronics enclosure 3602 contains a power entry module3610 and DC power conversion module 3612 for converting power from an ACsource (e.g., a wall outlet) and supplying DC power to the electronicssystem described in further detail below. According to some embodiments,power entry module 3610 provides a power connection configured toreceive mains electricity, for example, single-phase power from a walloutlet (e.g., a single-phase outlet providing approximately between 110and 120 volts at approximately 50-60 Hertz and rated at 15, 20 or 30amperes, approximately between 210 and 250 volts at approximately 50-60Hertz and rated at 15, 20 or 30 amperes, depending on the national orregional power standard delivered from wall outlets in the geographiclocation in which the MRI system is deployed).

Power entry module 3610 may be adapted to filter received mainselectricity delivered in accordance with the corresponding powerstandard so that the AC power is suitable for input to DC power module3612. DC power module 3612 may comprise one or more power supplies toconvert AC power to DC power that can be distributed at voltage levelsneeded by the various electronics components of portable MRI system3600. DC power module 3612 may include, for example, one or morecommercial power supplies configured to receive AC power as an input andsupply DC power as an output. Commercial power supplies are availablethat are configured to receive a wide variety of AC power. For example,available power supplies are capable of receiving AC power in a rangefrom approximately 85V to approximately 265V in a frequency range from50-60 Hz and configured to convert the AC input to deliver approximately1600 W of DC power (e.g., 380V, 4.2 A DC power). The AC input range onsuch exemplary commercial power supplies makes it suitable for use withthe most common, if not all, mains electricity sources worldwide. Thus,according to some embodiments, the MRI system can be configured to beessentially agnostic to different wall power standards, allowing the MRIsystem to be operated from wall power comprehensively across differentregions and/or countries, requiring only a change to the plug typerequired by the particular outlet.

DC power module 3612 may include one or more AC-DC power supplies and/orone or more DC-DC power converters configured to deliver power to one ormore backplanes at levels required by the different electroniccomponents of the MRI system (e.g., power amplifiers, console,controllers, thermal management, on-board computer, etc.), as discussedin further detail below. Individual electronic components may furtherinclude one or more power regulators to transform power distributed bythe backplanes to desired levels needed for the respective electroniccomponents. It should be appreciated that by providing low powerelectronic components having power demands that do not exceed theavailable power provided by mains electricity, power circuitry fortransforming three-phase power to single-phase power can be eliminated,facilitating a simplified power entry module 3610, reducing the size,cost and complexity of the power circuitry of the MRI system.

The electronics system illustrated in FIG. 36 comprises backplanes 3616and 3618 coupled to DC power module 3612 and configured to distributeoperating power at desired levels to various hardware components of theelectronics system. Backplane 3616 is configured to provide power toGPAs 3629 (e.g., any of the low power, low noise GPAs described above inconnection with FIGS. 20-34), RFPAs 3627 (e.g., RFPAs using any of thetechniques described in connection with low power RFPA 3500 illustratedin FIG. 35) and gradient coil(s) 3660. Backplane 3616 therefore providesthe connections to deliver power to the power amplifiers (e.g., GPA 3629and RFPA 3627) from DC power module 3612 and provides the connections todeliver amplified power to the corresponding magnetics components (e.g.,gradient coils 3628 and RF coil(s) 3626) from the respective poweramplifiers. According to some embodiments, backplane 3616 has multipleinputs to receive power at different power levels for distribution tothe power amplifiers. According to some embodiments, backplane 3616 hasinputs to receive power from DC power module 3612 at +/−48V at 4 A,+/−15V at 50 A and +48V at 3 A for distribution to the power amplifiers.However, it should be appreciated that the above power inputs to thebackplane are merely exemplary, and the number of power inputs as wellas the voltage and amplitude levels of the power inputs will depend onthe specific design needs of a given implementation. According to someembodiments, backplane 3616 is a printed circuit board, allowing fordistribution of power using PCB power connectors, eliminating the needfor expensive, bulky and lossy cable bundles between the power sourceand the electronic components.

Backplane 3618 is configured to provide power to various controllersincluding main controller 3632, shim controller 3630 and fan controller3680, various electronic components such as analog-to-digital converter(ADC) circuitry 3634, preamplifiers 3640 and magnetics component such asshim coil(s) 3670. According to some embodiments, backplane 3618comprises an input to receive power from DC power module 3612 at +48V, 4A for distribution to the components connected to the backplane.According to some embodiments, backplane 3618 includes a DC-DC converterto convert the 48V from DC power module 3612 to 12V for distribution toone or more of the connected components (e.g., main controller 3632). Aswith backplane 3616, backplane 3618 may be a printed circuit board todistribute power without the cable bundles used in conventional MRIpower systems to connect the power source to the electronic components,which may be located at relatively long distances. In the embodimentillustrated in FIG. 36, backplanes 3616 and 3618 are connected viaconnector 3617 to allow communication between the backplanes andcomponents connected thereto. It should be appreciated that thecomponents connected to the backplanes may be designed as boards or“cards” configured to connect to slots in the respective backplanes.However, one or more connected components can be implemented in adifferent manner, as the aspects are not limited in this respect.

The use of backplane(s) (e.g., exemplary backplanes 3616 and 3618)provides a number of advantages. As discussed above, backplanes allowelectronic components (e.g., power amplifiers, computers, console,controllers, etc.) to be connected to the backplane using PCB connectors(e.g., slots) to eliminate the long cabling conventionally used toconnect the power source to the electronic components, thus reducing thesize, complexity, cost and power losses that accompanies conventionalcabling systems. In addition, because the magnetics components arelocated proximate the electronic components (e.g., located directlyabove enclosure 3602), any necessary cabling connecting the magnetics tothe backplanes will be significantly reduced in size from the cablesused in conventional MRI, which typically had to connect powercomponents and other electronic components to magnetics componentslocated in separate rooms. Given these short distances, cables such asribbon cables can be used to connect the backplanes to the magneticscomponents to facilitate compact, simple and power efficient connectionbetween electronic and magnetics components of the MRI system. Moreover,the use of backplanes allows electronic components, such as poweramplifiers (e.g., GPAs 3629 and RFPA 3627), to be removed and replacedwithout needing to disconnect the magnetics components from therespective backplane.

Electronics enclosure 3602 also contains RFPA(s) 3627 and GPAs 3629,which in addition to being significantly lower power due to thelow-field strengths involved in the low-field and very low-fieldregimes, may also incorporate one or more of the low power techniquesdiscussed herein (e.g., as discussed above in connection with the GPAsillustrated in FIGS. 20-34 and/or the RFPAs discussed above inconnection with FIG. 35. In some embodiments, RFPA 3627 and/or GPA 3629may comprise a plurality of amplification stages using FETs or othersuitable switching components. In embodiments that include multipleamplification stages for one or both of RFPA 3627 and GPA 3629, each ofthe amplification stages may be associated with electromagneticshielding configured to shield the stage from electromagneticinterference. In contrast to large amplifier designs commonly used withconventional MRI systems, which require the use of large shieldingstructures, power amplifiers (e.g., RFPAs 3627, GPAs 3629) designed inaccordance with some embodiments employ smaller and/or simplerelectromagnetic shielding structures, further reducing the size,complexity and cost of a portable MRI system.

Additionally, the low power amplifiers (e.g., GPAs 3629 and RFPAs 3627)and lower drive currents for the gradient and RF coils may also simplifythermal management of the MRI system. For example, the low powerelectronic and magnetics components, in accordance with someembodiments, may be cooled using an air-cooled thermal managementsystem. For example, low power MRI system 3600 includes a fan controller3680 to control one or more fans (e.g., fans 3682 a, 3682 b, 3682 c) toprovide air to cool power components of the system that are co-locatedin electronics enclosure 3602 and/or the magnetics components of thesystem. An enclosure 3606 for fan controller 3780 may be located outsideelectronics enclosure 3602, for example, adjacent to or integrated withthe housing for the magnetics components (see e.g., FIG. 37D) or,according to some embodiments, may be located within electronicsenclosure 3602. Conventional high power systems often requirewater-based cooling systems that not only increase the size, cost andcomplexity of the system, but require a water source to operate the MRIsystem. Eliminating the need for water-based cooling facilitatesportability of the MRI system because the thermal management system canbe operated from the electrical power source (e.g., mains electricity),removing the need for an external water source and removing watercirculation equipment from the MRI system.

Electronics enclosure 3602 also includes a main controller 3632 (e.g., aconsole) configured to provide control signals to drive the operation ofthe various other components (e.g., RF coils, gradient coils, etc.) ofthe portable MRI system to provide console control in real-time or nearreal-time. For example, main controller 3632 may be programmed toperform the actions described in connection controller 106 illustratedin FIG. 1. Many conventional MRI systems include a console controllerimplemented as a specialized high-performance computer to performsimilar functions. In some embodiments, main controller 3632 isimplemented using a field-programmable gate array (FPGA), which hassubstantially fewer power requirements compared to the high-performanceconsole computers used in conventional MRI systems, contributing to thereduction in cost, complexity and power consumption of a portable MRIsystem. In the embodiment illustrated in FIG. 36, main controller 3632is connected to computer 3614 (e.g., a personal computer grade processorand memory system) to communicate between the two components. Computer3614 may include its own power converter and power supply and maytherefore have a separate connection to 3615 to power entry module 3610.

Portable MRI system 3600 also includes pre-amplifiers 3640 located in apre-amplifier enclosure 3604 to receive signals from RF coil(s) 3626.Pre-amplifiers 3640 are coupled to analog-to-digital converter (ADC)circuitry 3634 located within electronics enclosure 3602. ADCs 3634receive analog signals from RF coils 3626 via pre-amplifier circuitry3640 and convert the analog signals to digital signals that can beprocessed by computer 3614, including by transmitting signals to anexternal computer, for example, via a wireless connection (e.g.,transmitting digital signals to a smartphone, tablet computer, notepad,etc. used by an operator to initiate and/or control the imagingprotocol). RF coil(s) 3626 may include one or more noise coils, one ormore RF receive coils configured to detect MR signals and/or one or moreRF coils that operate as both noise coils and RF receive coils.Accordingly, signals received from RF coil(s) 3626 may include signalsrepresenting electromagnetic noise and/or signals representing MR data.Techniques for utilizing these signals in a noise reduction system tofacilitate operation of the portable MRI system outside of speciallyshielded rooms are discussed in further detail below (e.g., inconnection with FIGS. 41A-D and 42). Enclosure 3604 for preamplifiers3640 may be located outside electronics enclosure 3602, for example,adjacent to or integrated with the housing for the magnetics components(see e.g., FIG. 37D) or, according to some embodiments, may be locatedwithin electronics enclosure 3602.

As discussed above, MR data received from coils(s) 3626 may be processedby computer 3614 to suppress noise or otherwise prepare the MR data forimage reconstruction. According to some embodiments, MR data istransmitted to one or more external computers to perform imagereconstruction (e.g., MR data may be transmitted wirelessly to a mobiledevice and onto to secure server(s) in the cloud, or MR data may betransmitted directly to one or more servers for further processing).Alternatively, image reconstruction may be performed by computer 3614.The inventors have recognized that off-loading computation intensiveprocessing (e.g., image reconstruction and the like) to one or moreexternal computers reduces the power consumption of the on-boardcomputer 3614 and eliminates the need to use an on-board computer withsignificant processing power, reducing the cost and power consumption ofsuch implementations.

Electronics enclosure 3602 also provides containment for shim controller3630 configured to control the operation of one or more shim coil(s)3670 to improve the field homogeneity in an imaging field of view. Dueto the lower output currents required to control the operation of shimcoil(s) 3670 in a low-field MRI system, the electronics used toimplement shim controller 3630 may be smaller and/or simpler in similarways described above for RFPA 3627 and GPA 3629. For example, simplelow-power switches may be used to reduce the size and complexity of theshim controller, thereby facilitating the implementation of a portableMRI system. As discussed above, electronics enclosure 3602 may form, inpart, the base of portable MRI system 3600, the base supporting themagnetics components of the MRI system. For example, electronicsenclosure 3602 may form, in part, a base similar to base 1950 ofportable MRI system 1900 illustrated in FIGS. 19A and 19B. Accordingly,the components of an MRI system can be co-located on or within astandalone unit to provide portable MRI system 3600.

FIGS. 37A-D illustrate an exemplary arrangement of components of aportable MRI system. In particular, FIG. 37A illustrates a circularhousing 3702 that forms a part of base 3750 of portable MRI system 3700(shown FIG. 37D). Housing 3702 may house the components described inconnection with electronics enclosure 3602 illustrated in FIG. 36.Housing 3702 comprises a chassis or frame 3755 configured to secure theelectronic components and provide support for the magnetics componentspositioned on top of base 3750, as shown in FIG. 37D. Frame 3755separates housing 3702 into a number of partitions, including partition3702A that houses a first backplane to connect the power source to thepower amplifiers and to connect the power amplifiers to thecorresponding magnetics components, and partition 3702B that houses asecond backplane to connect the power source to the various controllers(e.g., a computer, main controller/console, shim controller, etc.) andto ADCs for digitizing signals from the RF coils. Partitions 3702A and3702B may, for example, house the electronic components illustrated inenclosure 3602 in FIG. 36, except for the power entry which is locatedin partition 3702C in FIG. 37A. Additionally, housing 3702 includes apartition 3702D for a motor that provides a power assist to facilitatetransporting or moving the portable MRI system to different locations.For example, portable MRI system 3700 may include one or more motorizedwheels that can be engaged when moving portable MRI system 3700 todifferent locations, as discussed in further detail below in connectionwith FIGS. 39A and 39B. Housing 3700 is manufactured having a diameterD, which may be chosen to facilitate moving the portable MRI system intypical spaces where the MRI system may be utilized (e.g., in emergencyrooms, intensive care units, operating rooms, etc.). According to someembodiments, housing 3700 has a diameter in a range between 25 and 40inches. For example, exemplary housing 3700 may have a diameter ofapproximately 32 inches to allow for relative ease in maneuvering thesystem in spaces where the portable MRI system is intended to beoperated.

FIGS. 37B and 37C illustrate different views of circular housing 3702 aspart of a base 3750 of portable MRI system 3700. The views in FIGS. 37Band 37C show the arrangement of electronic components within partitions3702A-C formed by frame 3755, with the backplanes located betweenpartitions 3702A and 3702B. FIG. 37D illustrates a portable MRI device3700 showing the magnetics components arranged atop based 3750. Inparticular, magnets 3722 a and 3722 b form, at least in part, a B₀magnet and gradient coils 3728 a and 3728 b provide X-gradient,Y-gradient and Z-gradient coils for portable MRI system 3700. As shown,portable MRI system may have a maximum horizontal width W thatfacilitates the maneuverability of the system within the facilities inwhich the MRI system is used. According to some embodiments, the maximumhorizontal dimension of a portable MRI system is in a range between 40and 60 inches and, more preferably, in a range between 35 and 45 inches.For example, exemplary portable MRI system 3700 has a maximum horizontalwidth of approximately 40 inches.

FIGS. 38A-F illustrate a number of exemplary steps in constructing aportable MRI system 3800. In FIG. 38A, a B₀ magnet 3810 comprising upperpermanent magnet 3810 a, lower permanent magnet 3810 b and yoke 3820 ismounted atop a base 3850, a portion of which is illustrated in FIG. 38A(the full base 3850 is illustrated in FIG. 38F). The upper and lowerpermanent magnets 3810 a and 3810 b are formed from a plurality ofconcentric rings of permanent magnet blocks, for example, similar to thepermanent magnets rings described in connection with FIGS. 16-18, thoughany configuration of permanent magnet rings may be used. B₀ magnet 3810and yoke 3820 may be constructed to be relatively light weight, forexample, using the techniques and materials described above inconnection with FIGS. 3-18 so that the total weight of the completedportable MRI system 3800, as shown in FIG. 38F, weighs less than 1,500hundred pounds and, more preferably, less than 1000 pounds. Accordingly,portable MRI system 3800 may be transported to different locations bypersonnel, with or without motor assist capabilities, examples of whichare described in further detail below.

B₀ magnet 3810 may be configured to produce a B₀ magnetic field in thevery low field strength regime (e.g., less than or equal toapproximately 0.1 T). For example, portable MRI system 3800 may beconfigured to operate at a magnetic field strength of approximately 64mT, though any low-field strength may be used. B₀ magnetic fieldstrengths in the very low-field regime facilitate a 5-Gauss line (e.g.,the perimeter outside of which the fringe magnetic field from the B₀magnet is 5 Gauss or less) that remains close to the portable MRIsystem. For example, according to some embodiments, the 5-Gauss line hasa maximum dimension of less than seven feet and, more preferably, lessthan 5 feet and, even more preferably, less than 4 feet. In addition tousing very low field strengths, shielding may be provided to reduce thevolume of the region inside the 5-Gauss line, as discussed in furtherdetail below.

As shown in FIG. 38A, provided on top of one or more of the permanentmagnet rings are permanent magnet shims 3830 configured to improve theprofile of the B₀ magnetic field produced by B₀ magnet 3810. Asdiscussed above, one exemplary technique for addressing the relativelylow SNR characteristic of the low-field regime is to improve thehomogeneity of the B₀ field by the B₀ magnet. In general, a B₀ magnetrequires some level of shimming to produce a B₀ magnetic field with aprofile (e.g., a B₀ magnetic field at the desired field strength and/orhomogeneity) satisfactory for use in MRI. In particular, productionfactors such as design, manufacturing tolerances, imprecise productionprocesses, environment, etc., give rise to field variation that producesa B₀ field having unsatisfactory profile after assembly/manufacture. Forexample, after production, exemplary B₀ magnets 200, 300 and/or 1600described above may produce a B₀ field with an unsatisfactory profile(e.g., inhomogeneity in the B₀ field unsuitable for imaging) that needsto be improved or otherwise corrected, typically by shimming, to produceclinically useful images.

Shimming refers to any of various techniques for adjusting, correctingand/or improving a magnetic field, often the B₀ magnetic field of amagnetic resonance imaging device. Similarly, a shim refers to something(e.g., an object, component, device, system or combination thereof) thatperforms shimming (e.g., by producing a magnetic field). Techniques forfacilitating more efficient and/or cost effective shimming for a B₀magnet for MRI are described in U.S. application Ser. No. 15/466,500('500 application), titled “Methods and Apparatus for Magnetic FieldShimming,” and filed on Mar. 22, 2017, which is herein incorporated byreference in its entirety.

Exemplary permanent magnet shims 3830 a, 3830 b, 3830 c and 3830 d maybe provided, for example, using any of the shimming techniques describedin the '500 application. In particular, the configuration or pattern(e.g., shape and size) of permanent magnet shims 3830 a-d may bedetermined by computing a magnetic field correction and determining amagnetic pattern for the permanent magnet shims to provide, at least inpart, the magnetic field correction. For example, permanent magnet shims3830 a-d may compensate for effects on the B₀ magnetic field resultingfrom asymmetric yoke 3820. For example, the pattern of the permanentmagnet shims 3830 a-d may be determined to mitigate and/or substantiallyeliminate non-uniformity in the B₀ magnetic field resulting from theeffects of yoke 3820 and/or more compensate for other non-uniformitiesin the B₀ magnetic field resulting from, for example, imperfectmanufacturing processes and materials to improve the profile (e.g.,strength and/or homogeneity) of the B₀ magnet. It should be appreciatedthat in the embodiment illustrated in FIG. 38A, permanent magnetic 3810a also has permanent magnet shims provided thereon that are not visiblein the view illustrated in FIG. 38A.

FIGS. 38B and 38C illustrate a vibration mount for the gradient coils ofportable MRI system 3800. As illustrated in FIG. 38B, vibration mount3840 includes portions positioned over the outer permanent magnet ringand fastened into place. In particular, circular arc segments 3842, ofwhich exemplary circular arc segments 3842A and 3842B are labeled, areaffixed to the frame on the outside of the outer permanent magnet ringand corresponding circular arc segments 3844, of which exemplarycircular arc segments 3844A and 3844B are labeled, are affixed to theframe on the inside of the outer permanent magnet ring. Slats 3845, ofwhich exemplary slats 3845A-D are labeled, are fastened to the circulararc segments 3842 and 3844 to form a vibration mount on which thegradient coils are mounted, as illustrated in FIG. 38D. As shown in FIG.38C, additional circular arc segments 3846 and 3848 are arranged betweenthe inner permanent magnet rings to facilitate fastening the gradientcoils to vibration mount 3840. FIG. 38C illustrates a completedvibration mount 3840 configured so that the gradient coils (e.g., alaminate panel on which gradient coils are fabricated) can be fastenedto the frame of the B₀ magnet to provide spacing between the gradientcoils and the permanent magnet shims and rings of the B₀ magnet 3810,and to provide vibration damping to reduce the acoustic noise andvibration of the gradient coils during operation. It should beappreciated that in the embodiment illustrated in FIGS. 38B-C, avibration mount is also provided on the upper permanent magnet that isnot visible in the view illustrated in FIGS. 38B and 38C.

FIG. 38D illustrates a laminate panel 3828 having gradient coilsfabricated thereon fastened to vibration mount 3840. For example,laminate panel 3828 may have one or more x-gradient coils, one or morey-gradient coils and/or one or more z-gradient coils patterned into oneor more layers of laminate panel 3828. One or more other magneticscomponents may also be fabricated on laminate panel 3828, such as one ormore shim or correction coils for the B₀ magnet 3810. Techniques forfabricating magnetics components on laminate panels is described in U.S.Pat. No. 9,541,616 ('616 patent), titled “Low-Field Magnetic ResonanceImaging Methods and Apparatus,” issued Jan. 10, 2017, which is hereinincorporated by reference in its entirety. It should be appreciated thatin the embodiment illustrated in FIG. 38D, a laminate panel comprisingone or more gradient coils (e.g., gradient coils for the X, Y and Zdirections) is also fastened to the vibration mount provided on theupper permanent magnet that is not visible in the view illustrate inFIG. 38D to provide the gradient magnetic fields needed for MRI.

FIG. 38E illustrates additional permanent magnet shims 3830′ affixedover the laminate panel 3828 illustrated in FIG. 38D. Permanent magnetshim 3830′ may provide fine shimming for the B₀ magnet. In particular,using any of the techniques described in the '500 applicationincorporated herein, the magnetic pattern of permanent magnet shim 3830′may be determined by computing a magnetic field correction anddetermining a magnetic pattern for the permanent magnet shim to provide,at least in part, the magnetic field correction. The patterned permanentmagnet shim 3830 may be affixed to a substrate 3832 so that it can besecured to the portable MRI system on top of the laminate panel (e.g.,using any of the techniques for patterning described in the '500application). In this manner, permanent magnet shims 3830 illustrated inFIG. 38A may provide a coarse shimming and permanent magnet shim 3830′may provide a finer shim to improve the profile of the B₀ magnetic fieldproduced by B₀ magnet 3810 (e.g., to correct for a B₀ offset and/or toimprove the homogeneity of the B₀ magnetic field). It should beappreciated that in the embodiment illustrated in FIG. 38E, anotherpermanent magnet shim may be affixed to the frame over the laminatepanel on the upper permanent magnet that is not visible in the viewshown in FIG. 38E to correct and/or improve the profile of the B₀magnetic field produced by permanent magnet 3810. The shims provided(e.g., permanent magnet shims 3830, 3830′ and/or shim coils fabricatedon the laminate panels along with the gradient coils) facilitates ahomogeneous B₀ magnetic field suitable for obtaining clinically usefulimages (e.g., the images illustrated in FIGS. 47-50 below).

FIG. 38F illustrates portable MRI system 3800 with housings or outercoverings over the magnetics components illustrated in FIGS. 38A-E. Inparticular, housing 3815A and 3815B provide covering for the B₀permanent magnet 3810, permanent magnet shims 3830 and 3830′, andlaminate panel 3828 comprising the gradient coils for the system for theupper and lower portions of the B₀ magnet, respectively. Housing 3825provides a covering for yoke 3828 and, according to some embodiments,houses preamplifiers (e.g., preamplifiers 3640 and 3740 illustrated inFIGS. 36 and 37D, respectively) and a fan controller (e.g., fancontroller 3680 and 3780 illustrated in FIGS. 36 and 37D, respectively)that controls the thermal management for the system. The magneticscomponents of portable MRI system 3800 are supported by base 3850comprising a housing 3802 for housing the electronic components of theportable MRI system (e.g., the electronic components discussed aboveconfigured to operate using mains electricity, such as from a standardwall outlet). Portable MRI system 3800 may be sized as discussed aboveto facilitate maneuverability of the portable MRI system 3800 so thatthe system can be brought to the patient. In addition, portable lowfield MRI system 3800 may be constructed of materials and designed to belight weight, preferably less than 1,500 pounds and, more preferably,less than 1,000 pounds.

As discussed above, a factor in developing a portable MRI system is theability to operate the MRI system in generally unshielded, partiallyshielded environments (e.g., outside of specially shielded rooms orencompassing cages or tents). To facilitate portable MRI that can beflexibly and widely deployed and that can be operated in differentenvironments (e.g., an emergency room, operating room, office, clinic,etc.), the inventors have developed noise reduction systems comprisingnoise suppression and/or avoidance techniques for use with MRI systemsin order to eliminate or mitigate unwanted electromagnetic noise, reduceits impact on the operation of the MRI systems and/or to avoid bands inthe electromagnetic spectra where significant noise is exhibited.

Performance of a flexible low-field MRI systems (e.g., a generallymobile, transportable or cartable system and/or a system that can beinstalled in a variety of settings such as in an emergency room, officeor clinic) may be particularly vulnerable to noise, such as RFinterference, to which many conventional high field MRI systems arelargely immune due to being installed in specialized rooms withextensive shielding. To facilitate low field MRI systems that can beflexibly and widely deployed, the inventors have developed noisereduction systems that employ one or more noise suppression techniquesfor use with low-field MRI systems in order to eliminate or mitigateunwanted noise or to reduce its impact on the operation of the low-fieldsystems.

According to some embodiments, noise suppression and/or avoidancetechniques are based on noise measurements obtained from theenvironment. The noise measurements are subsequently used to reduce thenoise present in MR signals detected by the low-field MRI system (e.g.,a system having a B₀ field of approximately 0.2 T or less, approximately0.1 T or less, approximately 50 mT or less, approximately 20 mT or less,approximately 10 mT or less, etc.) during operation, either bysuppressing the environmental noise, configuring the low-field MRIsystem to operate in a frequency band or bin having less noise, usingsignals obtained from multiple receive coils, or some combinationtherewith. Thus, the low-field MRI system compensates for noise presentin whatever environment the system is deployed and can therefore operatein unshielded or partially shielded environments so that MRI is notlimited to specialized shielded rooms.

Noise suppression techniques developed by the inventors, examples ofwhich are descried in further detail below, facilitate operation of MRIsystems outside shielded rooms and/or that have varying levels of devicelevel shielding of the imaging region of the system. Accordingly, MRIsystems employing one or more of the noise suppression techniquesdescribed herein may be employed where needed and in circumstances whereconventional MRI is unavailable (e.g., in emergency rooms, operatingrooms, intensive care units, etc.). While aspects of these noisesuppression techniques may be particularly beneficial in the low-fieldcontext where extensive shielding may be unavailable or otherwise notprovided, it should be appreciated that these techniques are alsosuitable in the high-field context and are not limited for use with anyparticular type of MRI system.

Using the techniques described herein, the inventors have developedportable, low power MRI systems capable of being brought to the patient,providing affordable and widely deployable MRI where it is needed. FIGS.39A and 39B illustrate views of a portable MRI system, in accordancewith some embodiments. Portable MRI system 3900 comprises a B₀ magnet3910 formed in part by an upper magnet 3910 a and a lower magnet 3910 bhaving a yoke 3920 coupled thereto to increase the flux density withinthe imaging region. The B₀ magnet 3910 may be housed in magnet housing3912 along with gradient coils 3915 (e.g., any of the gradient coilsdescribed in U.S. application Ser. No. 14/845,652, titled “Low FieldMagnetic Resonance Imaging Methods and Apparatus” and filed on Sep. 4,2015, which is herein incorporated by reference in its entirety).According to some embodiments, B₀ magnet 3910 comprises anelectromagnet, for example, an electromagnet similar to or the same aselectromagnet 210 illustrated in FIG. 2. According to some embodiments,B₀ magnet 3910 comprises a permanent magnet, for example, a permanentmagnet similar to or the same as permanent magnet 300 illustrated inFIG. 3A or permanent magnet 1600 illustrated in FIG. 16.

Portable MRI system 3900 further comprises a base 3950 housing theelectronics needed to operate the MRI system. For example, base 3950 mayhouse the electronics discussed above in connection with FIGS. 36-38,including power components configured to operate the MRI system usingmains electricity (e.g., via a connection to a standard wall outletand/or a large appliance outlet). For example, base 3970 may house lowpower components, such as those described herein, enabling at least inpart the portable MRI system to be powered from readily available walloutlets. Accordingly, portable MRI system 3900 can be brought to thepatient and plugged into a wall outlet in the vicinity.

Portable MRI system 3900 further comprises moveable slides 3960 that canbe opened and closed and positioned in a variety of configurations.Slides 3960 include electromagnetic shielding 3965, which can be madefrom any suitable conductive or magnetic material, to form a moveableshield to attenuate electromagnetic noise in the operating environmentof the portable MRI system to shield the imaging region from at leastsome electromagnetic noise. As used herein, the term electromagneticshielding refers to conductive or magnetic material configured toattenuate the electromagnetic field in a spectrum of interest andpositioned or arranged to shield a space, object and/or component ofinterest. In the context of an MRI system, electromagnetic shielding maybe used to shield electronic components (e.g., power components, cables,etc.) of the MRI system, to shield the imaging region (e.g., the fieldof view) of the MRI system, or both.

The degree of attenuation achieved from electromagnetic shieldingdepends on a number of factors including the type material used, thematerial thickness, the frequency spectrum for which electromagneticshielding is desired or required, the size and shape of apertures in theelectromagnetic shielding (e.g., the size of the spaces in a conductivemesh, the size of unshielded portions or gaps in the shielding, etc.)and/or the orientation of apertures relative to an incidentelectromagnetic field. Thus, electromagnetic shielding refers generallyto any conductive or magnetic barrier that acts to attenuate at leastsome electromagnetic radiation and that is positioned to at leastpartially shield a given space, object or component by attenuating theat least some electromagnetic radiation.

It should be appreciated that the frequency spectrum for which shielding(attenuation of an electromagnetic field) is desired may differdepending on what is being shielded. For example, electromagneticshielding for certain electronic components may be configured toattenuate different frequencies than electromagnetic shielding for theimaging region of the MRI system. Regarding the imaging region, thespectrum of interest includes frequencies which influence, impact and/ordegrade the ability of the MRI system to excite and detect an MRresponse. In general, the spectrum of interest for the imaging region ofan MRI system correspond to the frequencies about the nominal operatingfrequency (i.e., the Larmor frequency) at a given B₀ magnetic fieldstrength for which the receive system is configured to or capable ofdetecting. This spectrum is referred to herein as the operating spectrumfor the MRI system. Thus, electromagnetic shielding that providesshielding for the operating spectrum refers to conductive or magneticmaterial arranged or positioned to attenuate frequencies at least withinthe operating spectrum for at least a portion of an imaging region ofthe MRI system.

In portable MRI system 3900 illustrated in the moveable shields are thusconfigurable to provide shielding in different arrangements, which canbe adjusted as needed to accommodate a patient, provide access to apatient and/or in accordance with a given imaging protocol. For example,for the imaging procedure illustrated in FIG. 40A (e.g., a brain scan),once the patient has been positioned, slides 4060 can be closed, forexample, using handle 4062 to provide electromagnetic shielding 4065around the imaging region except for the opening that accommodates thepatient's upper torso. In the imaging procedure illustrated in FIG. 40B(e.g., a scan of the knee), slides 4060 may be arranged to have openingson both sides to accommodate the patient's legs. Accordingly, moveableshields allow the shielding to be configured in arrangements suitablefor the imaging procedure and to facilitate positioning the patientappropriately within the imaging region.

As discussed above, a noise reduction system comprising one or morenoise reduction and/or compensation techniques may also be performed tosuppress at least some of the electromagnetic noise that is not blockedor sufficiently attenuated by shielding 3965. In particular, asdiscussed in the foregoing, the inventors have developed noise reductionsystems configured to suppress, avoid and/or reject electromagneticnoise in the operating environment in which the MRI system is located.According to some embodiments, these noise suppression techniques workin conjunction with the moveable shields to facilitate operation in thevarious shielding configurations in which the slides may be arranged.For example, when slides 4060 are arranged as illustrated in FIG. 40B,increased levels of electromagnetic noise will likely enter the imagingregion via the openings. As a result, the noise suppression componentwill detect increased electromagnetic noise levels and adapt the noisesuppression and/or avoidance response accordingly. Due to the dynamicnature of the noise suppression and/or avoidance techniques developed bythe inventors, the noise reduction system is configured to be responsiveto changing noise conditions, including those resulting from differentarrangements of the moveable shields. Thus, a noise reduction system inaccordance with some embodiments may be configured to operate in concertwith the moveable shields to suppress electromagnetic noise in theoperating environment of the MRI system in any of the shieldingconfigurations that may be utilized, including configurations that aresubstantially without shielding (e.g., configurations without moveableshields), as discussed in further detail below.

To ensure that the moveable shields provide shielding regardless of thearrangements in which the slides are placed, electrical gaskets may bearranged to provide continuous shielding along the periphery of themoveable shield. For example, as shown in FIG. 39B, electrical gaskets3967 a and 3967 b (see also FIG. 45C) may be provided at the interfacebetween slides 3960 and magnet housing to maintain to provide continuousshielding along this interface. According to some embodiments, theelectrical gaskets are beryllium fingers or beryllium-copper fingers, orthe like (e.g., aluminum gaskets), that maintain electrical connectionbetween shields 3965 and ground during and after slides 3960 are movedto desired positions about the imaging region. According to someembodiments, electrical gaskets 3967 c are provided at the interfacebetween slides 3960, as illustrated in FIG. 40B, so that continuousshielding is provided between slides in arrangements in which the slidesare brought together. Accordingly, moveable slides 3960 can provideconfigurable shielding for the portable MRI system.

To facilitate transportation, a motorized component 3980 is provide toallow portable MRI system to be driven from location to location, forexample, using a control such as a joystick or other control mechanismprovided on or remote from the MRI system. In this manner, portable MRIsystem 3900 can be transported to the patient and maneuvered to thebedside to perform imaging, as illustrated in FIGS. 40A and 40B. Asdiscussed above, FIG. 40A illustrates a portable MRI system 4000 thathas been transported to a patient's bedside to perform a brain scan.FIG. 40B illustrates portable MRI system 4000 that has been transportedto a patient's bedside to perform a scan of the patient's knee.

The portable MRI systems described herein (e.g., MRI systems illustratedin FIGS. 19 and 39-40) may be operated from a portable electronicdevice, such as a notepad, tablet, smartphone, etc. For example, tabletcomputer 3975 may be used to operate portable MRI system to run desiredimaging protocols and to view the resulting images. Tablet computer maybe connected to a secure cloud to transfer images for data sharing,telemedicine and/or deep learning on the data sets. Any of thetechniques of utilizing network connectivity described in U.S.application Ser. No. 14/846,158, titled “Automatic Configuration of aLow Field Magnetic Resonance Imaging System,” filed Sep. 4, 2015, whichis herein incorporated by reference in its entirety, may be utilized inconnection with the portable MRI systems described herein.

FIG. 39C illustrates another example of a portable MRI system, inaccordance with some embodiments. Portable MRI system 4000 may besimilar in many respects to portable MRI systems illustrated in FIGS.16, 39A and 39B. However, slide 4060 are constructed differently, as isshielding 3965′, resulting in electromagnetic shields that are easierand less expensive to manufacture. As discussed above, a noise reductionsystem may be used to allow operation of a portable MRI system inunshielded rooms and with varying degrees of shielding about the imagingregion on the system itself, including no, or substantially no,device-level electromagnetic shields for the imaging region, asdiscussed in further detail below (e.g., in connection with FIGS. 41A-Dand 42).

It should be appreciated that the electromagnetic shields illustrated inFIGS. 39-40 are exemplary and providing shielding for an MRI system isnot limited to the example electromagnetic shielding described herein.Electromagnetic shielding can be implemented in any suitable way usingany suitable materials. For example, electromagnetic shielding may beformed using conductive meshes, fabrics, etc. that can provide amoveable “curtain” to shield the imaging region. Electromagneticshielding may be formed using one or more conductive straps (e.g., oneor more strips of conducting material) coupled to the MRI system aseither a fixed, moveable or configurable component to shield the imagingregion from electromagnetic interference, some examples of which aredescribed in further detail below. Electromagnetic shielding may beprovided by embedding materials in doors, slides, or any moveable orfixed portion of the housing. Electromagnetic shields may be deployed asfixed or moveable components, as the aspects are not limited in thisrespect.

Accordingly, aspects of the technology described herein relate toimproving the performance of low-field MRI systems in environments wherethe presence of noise, such as RF interference, may adversely impact theperformance of such systems. In some embodiments, a low-field MRI systemmay be configured to detect noise (e.g., environmental electromagneticnoise, internal system noise, radio frequency interference, etc.) and,in response, adapt the low-field MRI system to reduce the impact of thenoise on the operation of the system. The low-field MRI system may beconfigured to reduce the impact of the noise by suppressing noise in theRF signal obtained by the RF receive coil, by generating RF signals thatdestructively interfere with noise in the environment (e.g., RFinterference), by adjusting characteristics of the magnetic fieldsproduced (e.g., adjusting the magnetic field strength of the B0 magnet)and/or received by the low-field MRI system so that the transmit/receivecoils operate in a frequency band satisfactorily free from interference,or using a combination of these techniques.

According to some embodiments, noise suppression techniques describedherein allow a MRI system to be operated in unshielded or partiallyshielded environments and/or with or without device level shielding ofthe imaging region (e.g., shielding provided on the low-field MRI deviceitself to shield the imaging region from electromagnetic interference),at least in part by adapting noise compensation to the particularenvironment in which the MRI system is deployed. As a result, deploymentof an MRI system is not confined to specially shielded rooms or othercustomized facilities and instead can be operated in a wide variety ofenvironments.

In some embodiments, a system may be configured to obtain informationabout noise in the system's environment or within the system itself(e.g., RF interference) and suppress noise in the RF signal measured bythe RF receive coil based, at least in part, on the obtainedinformation. The system may be configured to obtain information aboutnoise in the environment by using one or more auxiliary sensors. Theterm “auxiliary” is used to differentiate between a sensor or detectorcapable of detecting noise and the primary receive channel that receivesMR signals for use in MRI. It should be appreciated that, in someembodiments, an auxiliary sensor may also receive one or more MRsignals. For example, the low-field MRI system may comprise one or moreauxiliary RF receive coils positioned proximate to the primarytransmit/receive coil(s), but outside of the field of view of the B0field, to detect RF noise without detecting MR signals emitted by asubject being imaged. The noise detected by the auxiliary RF coil(s) maybe used to suppress the noise in the MR signal obtained by the primaryRF coil of the MRI system.

Such an arrangement has the ability to dynamically detect and suppressRF noise to facilitate the provision of, for example, a generallytransportable and/or cartable low-field MRI system that is likely to besubjected to different and/or varying levels of RF noise depending onthe environment in which the low-field MRI system is operated. That is,because noise suppression is based on the current noise environment,techniques described herein provide noise suppression capabilityspecific to the particular environment in which the system is deployed.The simplistic approach of subtracting samples of noise obtained by oneor more auxiliary sensors from the signal measured by the primaryreceive coil(s) generally provides unsatisfactory noise suppression,even if the gain of the noise detected by the auxiliary sensor(s) isadjusted. The primary receive coil(s) and the auxiliary sensor(s) maymeasure different noise signals because the primary coil(s) and theauxiliary sensor(s) may be in different locations, have differentorientations, and/or may have different physical characteristics (e.g.,may have a different number of coil turns, may differ in size, shape,impedance, or may be a different type of sensor altogether).

Different locations and/or orientations of the primary coil(s) and theauxiliary sensor(s) may lead to differences in the characteristics ofthe noise signals received by the primary coil and the auxiliary sensor.Different physical characteristics between the primary coil(s) andauxiliary sensor(s) may lead to frequency-dependent differences betweennoise signals received by the primary coil(s) and auxiliary sensor(s).As a result, subtracting the noise signal measured by one or moreauxiliary sensors from the signal measured by the primary coil(s) maynot adequately suppress noise detected by the primary coil(s). Even ifthe noise signal measured by the auxiliary sensor(s) were scaled by aconstant in an attempt to compensate for differences in the gain of thenoise signals received by the primary coil(s) and auxiliary sensor(s),such compensation would not account for frequency-dependent differencesin the noise signals.

Some noise suppression techniques employ a transform to suppress noisein the RF signal received by one or more primary receive coil(s) of alow-field MRI system. According to some embodiments, the transformoperates to transform a noise signal received via one or multipleauxiliary sensors (e.g., one or more auxiliary RF coils and/or othertypes of sensors described herein) to an estimate of the noise receivedby the primary receive coil (or multiple primary receive coils). In someembodiments, noise suppression may comprise: (1) obtaining samples ofnoise by using the one or more auxiliary sensor(s); (2) obtainingsamples of the MR data using the primary RF coil; (3) determining atransform; (4) transforming the noise samples using the transform; and(5) subtracting the transformed noise samples from the obtained MR datato suppress and/or eliminate noise.

The transform may be estimated from multiple (e.g., at least ten, atleast 100, at least 1000, etc.) calibration measurements obtained usingthe auxiliary sensor(s) and primary coil(s). Multiple calibrationmeasurements allow for estimating the transform with high accuracy. Thetransform may be computed in the time domain, frequency domain or acombination of both. According to some embodiments, a transform may beestimated from the plurality of calibration measurements. Multiplecalibration measurements allow for estimating the amplitude and phase ofthe transform for a plurality of frequency bins across the frequencyspectrum for which the transform is defined. For example, whenprocessing signals using a K-point DFT (e.g., where K is an integerequal to 128, 256, 512, 1024 etc.), multiple measurements may allow forestimating the amplitude and phase of the transform for each of the Kfrequency bins.

In some embodiments, multiple auxiliary receive coils may be used asauxiliary sensors to suppress noise received by the primarytransmit/receive coil(s) of a low-field MRI system. For example, in someembodiments, a low-field MRI system may include multiple RF coilspositioned/configured to sense the MR signal emitted by the subjectbeing imaged (e.g., multiple “primary” coils) and/or multiple coilspositioned/configured to receive noise data, but to detect little or noMR signal (e.g., multiple “auxiliary” coils). Such an arrangementfacilitates detection and characterization of multiple noise sources tosuppress a variety of noise that may be present in a given environment.Multiple primary receive coils may also be used that factor into thenoise characterization techniques described herein, as well as beingused to accelerate image acquisition via parallel MR, or in othersuitable ways, as discussed in further detail below.

In some embodiments, multiple auxiliary sensors may be used to performnoise compensation when there are multiple sources of noise in theenvironment of the low-field MRI system. For example, one or moreauxiliary RF coils and/or one or more other types of sensors may be usedto obtain information about the noise environment resulting from noiseproduced by multiple sources, which information in turn may be used toprocess the RF signal received by the primary receive coil(s) in orderto compensate for the noise produced by multiple sources. For example,in some embodiments, a multichannel transform may be estimated fromcalibration measurements obtained using multiple auxiliary sensors andthe primary RF coil(s), as described in more detail below. Themultichannel transform may represent the relationships among the noisesignals captured by the primary RF coil(s) and each of the multipleauxiliary sensors. For example, the transform may capture correlationamong the noise signals received by the multiple auxiliary sensors. Thetransform may also capture correlation among the noise signals receiveby the multiple auxiliary sensors and the noise signal received by theprimary RF coil(s).

In some embodiments, multiple auxiliary sensors may be used to performnoise suppression by: (1) obtaining samples of noise by using multipleauxiliary sensors; (2) obtaining samples of the MR data using theprimary RF coil(s); (3) obtaining a multichannel transform; (4)transforming the noise samples using the multichannel transform; and (5)subtracting the transformed noise samples from the obtained MR data tosuppress and/or eliminate noise.

In some embodiments, the multichannel transform may be estimated frommultiple (e.g., at least ten, at least 100, at least 1000, etc.)calibration measurements. According to some embodiments, multiplecalibration measurements are used to estimate the amplitude and phase ofthe transform for a plurality of frequency bins across which themultichannel transform is defined. For example, when processing signalsusing a K-point DFT (e.g., where K is an integer equal to 128, 256, 512,1024 etc.), multiple calibration measurements may allow for estimatingthe amplitude and phase of the multichannel transform for each of the Kfrequency bins.

According to some embodiments, the MR signal detected by one or moreprimary receive coils may also be utilized to characterize the noise tosuppress or eliminate noise from the MR data. In particular, theinventors have recognized that by repeating MR data acquisitions usingthe same spatial encoding (e.g., by repeating a pulse sequence with thesame operating parameters for the gradient coils), the “redundant” dataacquired can be used to characterize the noise. For example, if a pulsesequence is repeated with the same spatial encoding multiple times, theMR data obtained should in theory be the same. Thus, the difference inthe signals acquired from multiple acquisitions using the same spatialencoding can be presumed to have resulted from noise. Accordingly,multiple signals obtained from using the same spatial encoding may bephase shifted and subtracted (or added) to obtain a measure of thenoise.

According to some embodiments, noise characterized in this manner can beused to compute a transform or included as a channel in a multi-channeltransform, as discussed in further detail below. Alternatively, noisecharacterized in this manner can be used alone or in combination withother techniques to suppress noise from acquired MR signals. Forexample, a noise estimate obtained based on multiple MR signals obtainedusing the same spatial encoding may be used to suppress noise withoutcomputing a transform, as other suitable techniques may be used.

According to some embodiments, one or more sensors (e.g., one or more RFcoils or other sensors capable of detecting electromagnetic fields) maybe used to assess the noise background in a spectrum of interest toassess which band within the spectrum is cleanest from a noiseperspective so that the transmit/receive coil(s) may be configured tooperate in the identified frequency band. Accordingly, in someembodiments, a low-field MRI system may be adapted by adjusting thetransmit/receive coil(s) to operate at a frequency band having lessinterference relative to other frequency bands in which thetransmit/receive coil(s) can be configured to operate. For example, oneor more auxiliary RF coils may be configured to monitor noise acrossmultiple frequency bands over which the primary RF coil could operateand, the primary RF coil may be configured to operate at the frequencyband having the least amount of noise, as determined by the measurementsobtained using the auxiliary RF coils. In particular, an auxiliary RFcoil may be a wideband RF coil configured to measure the noise level(e.g., noise floor) across a wide band of frequencies. Based on thenoise measured across a frequency band of interest, the primarytransmit/receive coil(s) (e.g., which may be a narrowband coil) may beconfigured to operate in a band determined to have less noise than otherfrequency bands. Alternatively, multiple sensors may be provided, eachmeasuring noise levels in a respective frequency band. The primarytransmit/receive coil(s) may then be configured to operate in thefrequency band determined to have the least amount of noise present.

A significant source of interference for a low-field MRI system may beone or more power lines (e.g., power cords) supplying power to thelow-field MRI system. Accordingly, in some embodiments, a low-field MRIsystem is configured to measure directly any interference due to thepower line(s) and use the measurements to suppress or cancel suchinterference. For example, in some embodiments, a low-field MRI systemmay include one or more sensors coupled to a power line of the system tomeasure any RF signals produced or carried by the power line, and themeasurements obtained by the sensor(s) may be used as part of the noisesuppression techniques described herein (e.g., to further characterizethe noise environment and facilitate estimation of a comprehensivetransform).

In some embodiments, a low-field MRI system may include an antennacapacitively coupled to one of the power lines of the system and may beconfigured to use measurements obtained by the antenna to suppress noisein the RF signal received by the primary RF coil of the low-field MRIsystem. Such an antenna may be of any suitable type and, for example,may comprise a thin metal sheet wrapped around the power line and/or oneor more capacitors coupled to the power line. A low-field MRI system mayinclude multiple such antenna to detect noise resulting from any desirednumber of power lines supplying power to the system (or that otherwiseimpact the system) including, for example, hot lines carryingsingle-phase, two-phase, or three-phase power. In some instances, alow-field MRI system may include such an antenna for a ground wire. Asanother example, a low-field MRI system may include a sensor inductivelycoupled to a power line or multiple respective power lines (e.g., by useof a toroid or any other suitable method) to measure RF signals carriedby the power line such that these measurements may be used to suppressnoise in the RF signal measured by the primary RF coil of the low-fieldMRI system.

In some embodiments, a sensor's measurements of interference due to apower line may be used to suppress noise in the RF signal measured bythe primary RF receive coil by estimating a transform between theprimary RF receive coil and the sensor. This may be done in any suitableway and, for example, may be done using the techniques described hereinfor estimating a transform between the primary RF receive coil and anauxiliary RF receive coil. For example, noise characterized in thismanner may be used to estimate a transform alone or may be a channel ina multi-channel transform. Noise characterized by a sensor coupled toone or more power lines may be utilized in other manners (e.g., useddirectly to suppress noise), as the aspects are not limited in thisrespect.

According to some embodiments, noise in the environment may be detectedby coupling one or more sensors to one or more electromagneticinterference (EMI) shields. For example, a sensor may be connectedinductively or capacitively between one or more EMI shields and groundto detect the EMI captured by the shield. Noise characterized in thismanner may be used to suppress or eliminate noise from MR signalsdetected by the primary receive coil(s). For example, noisecharacterized by coupling a sensor to one or more EMI shields may beused to estimate a transform alone, or may be used as a channel in amulti-channel transform. Noise characterized by a sensor coupled to oneor more EMI shields may be utilized in other manners, as the aspects arenot limited in this respect.

Referring again to FIG. 1, MRI system 100 includes controller 106 (alsoreferred to as a console) having control electronics to sendinstructions to and receive information from power management system110. Controller 106 may be configured to implement one or more pulsesequences, which are used to determine the instructions sent to powermanagement system 110 to operate the magnetic components 120 in adesired sequence. For example, in a low-field MRI system, controller 106may be configured to control power management system 110 to operate themagnetic components 120 in accordance with a balance steady-state freeprecession (bSSFP) pulse sequence, a low-field gradient echo pulsesequence, a low-field spin echo pulse sequence, a low-field inversionrecovery pulse sequence, and/or any other suitable pulse sequence.Controller 106 may be implemented as hardware, software, or any suitablecombination of hardware and software, as aspects of the disclosureprovided herein are not limited in this respect.

In some embodiments, controller 106 may be configured to implement apulse sequence by obtaining information about the pulse sequence frompulse sequences repository 108, which stores information for each of oneor more pulse sequences. Information stored by pulse sequencesrepository 108 for a particular pulse sequence may be any suitableinformation that allows controller 106 to implement the particular pulsesequence. For example, information stored in pulse sequences repository108 for a pulse sequence may include one or more parameters foroperating magnetics components 120 in accordance with the pulse sequence(e.g., parameters for operating the RF transmit and receive coils 126,parameters for operating gradient coils 128, etc.), one or moreparameters for operating power management system 110 in accordance withthe pulse sequence, one or more programs comprising instructions that,when executed by controller 106, cause controller 106 to control system100 to operate in accordance with the pulse sequence, and/or any othersuitable information. Information stored in pulse sequences repository108 may be stored on one or more non-transitory storage media.

As illustrated in FIG. 1, controller 106 also interacts with computingdevice 104 programmed to process received MR data. For example,computing device 104 may process received MR data to generate one ormore MR images using any suitable image reconstruction process(es).Controller 106 may provide information about one or more pulse sequencesto computing device 104 for the processing of data by the computingdevice. For example, controller 106 may provide information about one ormore pulse sequences to computing device 104 and the computing devicemay perform an image reconstruction process based, at least in part, onthe provided information.

Computing device 104 may be any electronic device that may processacquired MR data and generate one or more images of the subject beingimaged. In some embodiments, computing device 104 may be a fixedelectronic device such as a desktop computer, a server, a rack-mountedcomputer, or any other suitable fixed electronic device that may beconfigured to process MR data and generate one or more images of thesubject being imaged. Alternatively, according to some embodiments of alow-field MRI system, computing device 104 may be a portable device suchas a smart phone, a personal digital assistant, a laptop computer, atablet computer, or any other portable device that may be configured toprocess MR data and generate one or images of the subject being imaged.In some embodiments, computing device 104 may comprise multiplecomputing devices of any suitable type, as the aspects are not limitedin this respect. A user 102 may interact with workstation 104 to controlaspects of the low-field MR system 100 (e.g., program the system 100 tooperate in accordance with a particular pulse sequence, adjust one ormore parameters of the system 100, etc.) and/or view images obtained bythe low-field MR system 100.

FIG. 41A shows illustrative components of a portion of an example a MRIsystem that may be used for performing noise suppression, in accordancewith some embodiments of the technology described herein. For example,transmit/receive system 4100 may form at least part of thetransmit/receive equipment (e.g., transmit/receive coils 126, one ormore controllers, etc.) of a low-field MRI system, such as any of theexemplary systems described in the above incorporated co-filed patentapplications. Transmit/receive system 4100 is configured to detect MRsignals emitted from excited atoms of a subject 4104 being imaged, andto characterize noise in the environment to suppress or remove thecharacterized noise from the detected MR signals, as discussed infurther detail below.

As shown in FIG. 41A, transmit/receive system 4100 comprises a primaryRF receive coil 4102 configured to measure MR signals emitted by thesubject 4104 in response to an excitation pulse sequence (e.g., a pulsesequence selected from pulse sequence repository 108 and executed bycontroller 102). The excitation pulse sequence may be produced byprimary RF receive coil 4102 and/or by one or more other transmit RFcoils arranged proximate subject 4104 and configured to produce suitableMR pulse sequences when operated. Primary receive coil 4102 may be asingle coil or may be a plurality of coils, which, in the latter case,may be used to perform parallel MRI. Tuning circuitry 4108 facilitatesoperation of primary receive coil 4102 and signals detected by RFcoil(s) 4102 are provided to acquisition system 4110, which may amplifythe detected signals, digitize the detected signals, and/or perform anyother suitable type of processing.

Transmit/receive system 4100 also includes auxiliary sensor(s) 4106,which may include any number or type of sensor(s) configured to detector otherwise measure noise sources in the environment and/orenvironmental noise produced by the MRI system itself. The noisemeasured by auxiliary sensor(s) 4106 may be characterized and used tosuppress noise in the MR signal detected by primary RF coil(s) 4102using techniques described in further detail below. After acquisitionsystem 4110 processes the signals detected by RF coil(s) 4102 andauxiliary sensor(s) 4106, acquisition system 4110 may provide theprocessed signals to one or more other components of the MRI system forfurther processing (e.g., for use in forming one or more MR images ofsubject 4104). Acquisition system 4110 may comprise any suitablecircuitry and may comprise, for example, one or more controllers and/orprocessors configured to control the MRI system to perform noisesuppression in accordance with embodiments described herein. It shouldbe appreciated that components illustrated in FIG. 41A may be configuredto detect MR signals generated by a MRI system and, for example, the RFcoils may be similar or the same as those described in the aboveincorporated co-field applications, or may be any other suitable type ofcoil.

In some embodiments, auxiliary sensor(s) 4106 may include one or moreauxiliary coils 4206 configure to measure noise from one or more noisesources in the environment in which the MRI system is operating, asshown in FIG. 41B. In some instances, the auxiliary RF coil(s) 4206 maybe constructed to be substantially more sensitive to ambient noise thanto any noise generated by the coil itself. For example, the auxiliary RFcoil 4206 may have a sufficiently large aperture and/or a number ofturns such that the auxiliary coil is more sensitive to noise from theenvironment than to noise generated by the auxiliary coil itself. Insome embodiments, auxiliary RF coil(s) 4206 may have a larger apertureand/or a greater number of turns than primary RF coil(s) 4102. However,auxiliary RF coil(s) 4206 may be the same as primary RF coil in thisrespect and/or may differ from primary RF coil(s) 4102 in otherrespects, as the techniques described herein are not limited to anyparticular choice of coils. For example, in some embodiments, anauxiliary sensor of a different type is used in place of an RF coil typesensor, as discussed in further detail below.

In the illustrative embodiment of FIG. 41B, auxiliary RF coil(s) 4206is/are located a distance 4205 apart from primary RF coil 4102. Thedistance 4205 may be selected such that auxiliary coil(s) 4206 is/aresufficiently far away from the sample 4104 to avoid sensing MR signalsemitted by the sample during imaging, but otherwise arranged as close aspossible to the primary RF coil 4102 so that auxiliary coil(s) 4206detect noise similar to the noise detected by primary coil(s) 4102. Inthis manner, the noise from one or more noise sources measured byauxiliary coil(s) 4206 and characterized using techniques discussedherein (e.g., by using the detected noise to calculate, at least inpart, a transform that can be used to suppress and/or eliminate noisepresent on detected MR signals) may be representative of the noisedetected by primary coil(s) 4102. It should be appreciated thatauxiliary coil(s) 4206 need not be RF coils, but may be any type ofsensor capable of detecting or measuring noise in the environment thatmay impact the performance of the MRI system, as the techniquesdescribed herein are not limited for use with any particular type ofsensor.

According to some embodiments, auxiliary sensor(s) 4106 may include oneor more auxiliary sensors 4306 configure to measure noise by couplingsensor(s) to one or more components of the MRI system, as schematicallyshown in FIG. 41C. For example, auxiliary sensors 4306 may include oneor more sensors coupled to one or more components of the MRI system orotherwise arranged to detect noise produced by the MRI system. Asdiscussed above, power cables are frequently a source of noise that canhave a negative impact on the operation of the MRI system and, inparticular, may produce noise that is detected by the one or moreprimary coils. According to some embodiments, auxiliary sensor(s) 4306include one or more sensors coupled (e.g., capacitively or inductively)to one or more power cables of the system to detect noise producedtherefrom. The detected noise may be characterized and used to suppressnoise from detected MR signals, for example, by using the detected noiseto produce, at least in part, a transform that characterizes noisedetected by the auxiliary sensor(s) 4306, or by being directly appliedto detected MR signals.

As discussed above, the low-field regime may facilitate systems that canbe utilized in a wide variety of circumstances and/or that can begenerally transported from one location to another. As a result,low-field MRI systems will frequently operate outside of speciallyshielded rooms. Thus, some low-field MRI systems may utilize partialshielding of one or more components of the system to prevent at leastsome EMI from reaching the shielded components. The inventors haveappreciated that by coupling one or more sensors to one or more EMIshields (e.g., a Faraday cage of one or more components or the like) ofthe system, the noise absorbed by the one or more EMI shields can bemeasured, characterized and used to suppress and/or eliminate noise fromdetected MR signals. According to some embodiments, auxiliary sensor(s)4306 include one or more sensors coupled between one or more EMI shieldsand ground to measure noise absorbed by the EMI shield that can be usedto facilitate noise suppression. For example, the noise detected fromthe EMI shield may be used to compute, at least in part, a transformthat can be utilized in suppressing and/or eliminating noise fromdetected MR signals. It should be appreciated that auxiliary sensor(s)4306 may include any other type of sensor capable of detecting noise, asthe aspects are not limited in this respect.

According to some embodiments, auxiliary sensor(s) 4106 include theprimary coil(s) itself as illustrated in FIG. 41D, wherein the primaryRF coil(s) are labeled both as primary receive coil 4102 and auxiliarysensor 4406 for the system, as the primary RF coil(s) may perform bothroles in some circumstances. As discussed above, the inventors haverecognized that certain pulse sequences facilitate using the signalsacquired from the primary coil(s) to also suppress noise thereon. Apulse sequence refers generally to operating transmit coil(s) andgradient coil(s) in a prescribed sequence to induce an MR response. Byrepeating the same pulse sequence using the same spatial encoding,“redundant” MR signals can be obtained and used to estimate noisepresent in the MR signals.

To address the relatively low signal-to-noise ratio (SNR) of low-fieldMRI, pulse sequences have been utilized that repeat MR data acquisitionsusing the same spatial encoding (e.g., by repeating a pulse sequencewith the same operating parameters to drive the gradient coils in thesame manner). The MR signals obtained over multiple acquisitions areaveraged to increase the SNR. For example, a balanced steady-state freeprecession (bSSFP) pulse sequence may be used to rapidly obtain MR dataover multiple acquisitions, which acquisitions are then averagedtogether to increase the SNR. The term “average” is used herein todescribe any type of scheme for combining the signals, includingabsolute average (e.g., mean), weighted average, or any other techniquethat can be used to increase the SNR by combining MR data from multipleacquisitions. Because the bSSFP pulse sequence does not require waitingfor the net magnetization to realign with the B₀ field betweensuccessive MR data acquisitions (e.g., successive acquisitions may beobtained without needing to wait for the transverse magnetization vectorto decrease to 0), multiple acquisitions may be rapidly obtained.However, any pulse sequence can be used to perform multiple acquisitionsat the same location, as the aspects are not limited in this respect.

The inventors have appreciated that the MR data obtained during multipleacquisitions performed using the same spatial encoding may be used tosuppress and/or eliminate noise from the detected MR signal. Asdiscussed above, when multiple acquisitions are performed by repeatingthe pulse sequence with the same spatial encoding, the MR signalsobtained should be the same or nearly the same and the differences canbe attributed to noise. As such, phase shifting the MR signal obtainedover multiple acquisitions and computing the difference between thesignals provides a means for evaluating the noise corrupting the MRdata. The difference may be obtained by phase shifting and either addingor subtracting the phase shifted MR signals depending on the type ofpulse sequence utilized. For example, the bSSFP pulse sequence flips thepolarity of the pulse sequence on subsequent acquisitions so that thedifference may be computed by adding MR signals that have beenappropriately shifted in phase. However, MR signals obtained using otherpulse sequences that do not flip the polarity may be subtracted afterbeing appropriately phase shifted to obtain the difference betweenmultiple MR acquisitions. Because multiple acquisitions (e.g., 10, 20,50, 100, 150 or more) obtained using the same spatial encoding mayalready be performed (and averaged) in the low-field context to achievesufficiently large SNR, using one or more of the acquisitions to computea noise estimate will not substantially increase acquisition times, ifat all.

The computed noise (e.g., the difference between MR signals obtainedover multiple acquisitions with the same spatial encoding can be used tosuppress and/or eliminate the noise in the detected MR signal. Accordingto some embodiments, the noise computed according to the above describedtechnique may be used to, at least in part, determine a transform thatcan be used to suppress and/or eliminate noise in the manner discussedin further detail below. However, noise computed by determining thedifference between multiple MR acquisitions can be utilized in otherways to suppress and/or eliminate noise, as the aspects are not limitedin this respect. For example, noise computed based on determining thedifference between multiple MR acquisitions obtained from the samelocation may be directly applied to detected MR signals or applied afterfurther processing. It should be appreciated that the noise computed bycomparing multiple acquisitions obtained using the same spatial encodingcan be used to dynamically suppress and/or eliminate noise from thedetected MR signals. In this way, noise cancellation dynamically adaptsto changing noise conditions in the environment.

As discussed above, noise detected by one or more auxiliary sensors,some examples of which are described in the foregoing, may be used tocharacterize the noise from one or more noise sources and suppressand/or eliminate noise from detected MR signals. According to someembodiments, the noise detected by one or more auxiliary sensors is usedto determine a transform that can be used to transform detected noise toan approximation of the noise detected by the one or more primaryreceive coils. According to some embodiments, noise detected by one ormore auxiliary sensors is applied to detected MR signals to suppressnoise without using a transform.

As a non-limiting example, a noise suppression component (e.g.,acquisition system 4110 illustrated in FIGS. 41A-D) may suppress noisein a signal s_(pri)(t), detected by primary RF coil 4102, by using thesignal s_(aux)(t), detected by auxiliary sensor 4106, and aprimary-to-auxiliary sensor (PA) transform H_(PA)(ω) via the followingexpression:s _(comp)(t)=s _(pri)(t)−

⁻¹ {H _(PA)(ω)S _(aux)(ω)},  (1)where S_(aux)(ω) is the Fourier transform of s_(aux)(t),

⁻¹{ } is the inverse Fourier transform operator, and s_(comp)(t) is thenoise-suppressed signal. It should be appreciated that the noisecompensation calculation of Equation (1) may be implemented in any ofnumerous ways and, for example, may be implemented in the frequencydomain or in the time domain, as the noise suppression techniquesdescribed herein are not limited in this respect. Exemplary techniquesfor estimating a PA transform are described in more detail below.

FIG. 42 is a flowchart of an illustrative process 4501 for performingnoise suppression, in accordance with some embodiments of the technologydescribed herein, including a detailed description of a technique fordetermining an exemplary transform, first with respect to a transformbetween an auxiliary sensor and a primary receive coil, followed by adescription of a transform between multiple auxiliary sensors and aprimary receive coil (multi-channel transform). It should be appreciatedthat a single or multi-channel transform may be computed for any numberof receive coils so that noise cancellation in this respect can beperformed using any number and type of auxiliary sensor and any numberand type of receive coil. Process 4501 may be performed by components ofany suitable MRI system and, for example, may be performed by componentsof MRI system 100 described with reference to FIG. 1 and the associatedcomponents illustrated in FIGS. 41A-D.

Process 4501 begins at acts 4502 and 4504, where a MRI system obtains MRdata by using a primary RF coil (e.g., RF coil 4102) and obtains noisedata using one or more auxiliary sensors (e.g., one or more RF coils4206 and/or one or more other sensors 4106, 4306, 4406, etc.). Asdiscussed above, any number of auxiliary sensors of any type may be usedto characterize the noise in the environment of the MRI system. Toillustrate aspects of the noise suppression techniques, the case of aprimary RF coil and an auxiliary sensor is first considered. The primaryRF coil and auxiliary sensor may operate to obtain MR and noise datasubstantially simultaneously such that the noise data acquired by theauxiliary sensor may be used to suppress noise in the MR data acquiredby the primary RF coil.

The signal obtained by the primary RF coil may comprise both noise andan MR signal emitted by the sample being imaged. For example, ifs_(pri)(t) represents the total signal measured by the primary RF coil,then s_(pri)(t) may be expressed as:s _(pri)(t)=m _(pri)(t)+n _(pri)(t),where m_(pri)(t) and n_(pri)(t) represent the MR signal and noisecomponents of the total signal measured by the primary RF coil. Assumingthat the auxiliary sensor measures a negligible amount of MR signal (dueto the placement of the auxiliary sensor relative to the primary RF coiland the sample being imaged), the signal measured by the auxiliarysensor contains mostly ambient RF noise. For example, if s_(aux)(t)represents the total signal measured by the auxiliary sensor, thens_(aux)(t) may be expressed according to:s _(aux)(t)=n _(aux)(t),where n_(aux)(t) is noise measured by the auxiliary sensor.

As discussed above, the noise components of the signals measured by theprimary RF coil and auxiliary sensor may be different (e.g., n_(pri)(t)may be different from n_(aux)(t)) due to physical differences betweenthe primary coil and auxiliary sensor as well as differences in locationand orientation. However, the inventors have appreciated that arelationship between the noise signals measured by the primary coil andthe auxiliary sensor may be established since both measure noise fromone or more common sources. Such a relationship may be, in someembodiments, represented by a primary to auxiliary transform. Forexample, the relationship may be represented by a primary to auxiliarytransform H_(PA)(ω) as detailed below.

For example, in some embodiments, each of the noise signals n_(pri)(t)and n_(aux)(t) may contain noise from several independent sourcesincluding, but not limited to, noise from one or more sources in theenvironment of the low-field MRI system, noise generated by the primaryRF coil and/or the auxiliary sensor, and noise generated by one or moreother components of the MRI system (e.g., noise generated by tuningcircuitry, acquisition system, power cables, etc.). Thus, for example,the noise signals n_(pri)(t) and n_(aux)(t) may be expressed as:n _(pri)(t)=c _(pri)(t)+u _(pri)(t), andn _(aux)(t)=c _(aux)(t)+u _(aux)(t)≅c _(aux)(t),where c_(pri)(t) and c_(aux)(t) represent correlated noise (i.e., thesignals c_(pri)(t) and c_(aux)(t) are correlated) generated by one ormore common noise sources detected by the primary coil and the auxiliarysensor, respectively, and where u_(pri)(t) and u_(aux)(t) representuncorrelated noise detected by the primary coil and auxiliary sensors,respectively (e.g., noise generated by the primary coil and auxiliarysensor themselves). As described above, in some embodiments, theauxiliary sensor may be configured such that it is more sensitive tonoise from the environment than noise generated by the sensor itself.For example, the auxiliary sensor may be an auxiliary RF coil having asufficiently large aperture and/or number of turns. As such, c_(aux)(t)may be substantially larger than u_(aux)(t) so thatn_(aux)(t)≅C_(aux)(t).

Each of the noise signals c_(pri)(t) and c_(aux)(t) can be expressed inrelation to the common noise source(s) through a respective measurementtransform. For example, in the Fourier domain, the Fourier transformsC_(pri)(ω) and C_(aux)(ω) of noise signals c_(pri)(t) and c_(aux)(t) canbe expressed as:C _(pri)(ω)=H _(pri)(ω)C _(s)(ω)C _(aux)(ω)=H _(aux)(ω)C _(s)(ω)where C_(s)(ω) is the Fourier transform of a common noise source andH_(pri)(ω) and H_(aux)(ω) respectively represent the channel between thecommon noise source and the primary receive coil and auxiliary sensor.Combining the above equations yields:

C_(pri)(ω) = H_(PA)(ω)C_(aux)(ω), where${{H_{PA}(\omega)} = \frac{H_{pri}(\omega)}{H_{aux}(\omega)}},$is the primary-to-auxiliary transform.

Returning to the discussion of process 4501, after the MR and noisesignals are acquired at acts 4502 and 4504, process 4501 proceeds to act4506, where a primary-to-auxiliary (PA) transform is obtained. In someembodiments, the PA transform may have been previously estimated so thatobtaining the PA transform at act 4506 comprises accessing arepresentation of the PA transform (e.g., a frequency-domain or atime-domain representation of the PA transform). In other embodiments,obtaining the PA transform at act 4506 may comprise estimating and/orupdating the estimate of the transform. Techniques for estimating a PAtransform are described in more detail below.

Next, at act 4508, the noise data obtained at act 4504 and the PAtransform obtained at act 4506 may be used to suppress or cancel noisein the MR data obtained at act 4502. This may be done using Equation (1)described above, using any equivalent formulation of Equation (1) (e.g.,the entire calculation may be performed in the frequency domain), or inany other suitable way.

As described above, a primary-to-auxiliary transform may be used tosuppress noise in the MR data acquired by a primary RF coil in a MRIsystem such as a low-field MRI system. In some embodiments, theprimary-to-auxiliary transform may be estimated from calibrationmeasurements obtained by the primary RF coil and the auxiliary sensor.This may be done in any suitable way. For example, the PA transform maybe estimated from calibration measurements obtained when no MR signal ispresent or when the strength of the MR signal is small relative to thestrength of the noise detected by the primary RF coil. As anotherexample, the PA transform may be estimated from calibration measurementsobtained when an MR signal is present (e.g., during operation of the MRIsystem). Any suitable number of calibration measurements may be used(e.g., at least 100, 100-1000, at least 1000, etc.). When moremeasurements are used, the PA transform may be estimated at a higherresolution (e.g., at more frequency values) and/or with increasedfidelity with respect to the actual noise environment. The PA transformmay be estimated using a least-squares estimation technique or any othersuitable estimation technique, as the techniques described herein arenot limited to any particular computational method.

According to some embodiments, a PA transform comprises a PA transformthat is estimated from the calibration measurements. As one non-limitingexample, when the signal acquired by the primary coil at times {t_(k)}does not contain any MR signal or when the strength of the MR signal issmall relative to the strength of the noise detected by the primary RFcoil, then s_(pri)(t_(k))=n_(pri)(t_(k)), so that the discrete Fouriertransform of s_(pri)(t_(k)) is given by:S _(pri)(ω_(k))=C _(pri)(ω_(k))+U _(pri)(ω_(k)),where C_(pri)(ω_(k)) is the discrete Fourier transform of C_(pri)(t_(k))and U_(pri)(ω_(k)) is the discrete Fourier transform of u_(pri)(t_(k)).Since C_(pri)(ω_(k))=H_(PA)(ω_(k))S_(ref)(ω_(k)), the discrete Fouriertransform of the signal received at the primary coil may be representedas a function of the discrete Fourier transform of the signal receivedat the auxiliary sensor according to:S _(pri)(ω_(k))=H _(PA)(ω_(k))S _(aux)(ω_(k))+U _(pri)(ω_(k))  (2)

Equation (2) represents a set of independent equations, one for eachfrequency component, ω_(k). Since both U_(pri) and H_(PA) are unknown,it may not be possible to determine H_(PA) from a single calibrationmeasurement. If M calibration measurements (e.g., at least 10, at least100, at least 1000 calibration measurements) are made such that multipleexamples of S_(pri) and S_(aux) for each frequency component areobtained, then the PA transform can be determined despite the unknownU_(pri), via any suitable estimation technique, for example, via leastsquares estimation. This is so because multiple measurements may be usedto average out the uncorrelated noise. Given M calibration measurements,a least squares estimator for the PA transform may be obtained byconsidering the following matrix equation for each frequency componentω_(k),

${\begin{bmatrix}{S_{pri}( \omega_{k} )}_{1} \\\vdots \\{S_{pri}( \omega_{k} )}_{M}\end{bmatrix} = {{H_{PA}( \omega_{k} )}\begin{bmatrix}{S_{aux}( \omega_{k} )}_{1} \\\vdots \\{S_{aux}( \omega_{k} )}_{M}\end{bmatrix}}},$which can be solved according to:

${H_{PA}( \omega_{k} )} = {{{\{ {\begin{bmatrix}{S_{aux}( \omega_{k} )}_{1} \\\vdots \\{S_{aux}( \omega_{k} )}_{M}\end{bmatrix}^{T}\begin{bmatrix}{S_{aux}( \omega_{k} )}_{1} \\\vdots \\{S_{aux}( \omega_{k} )}_{M}\end{bmatrix}} \}^{- 1}\begin{bmatrix}{S_{aux}( \omega_{k} )}_{1} \\\vdots \\{S_{aux}( \omega_{k} )}_{M}\end{bmatrix}}^{T}\begin{bmatrix}{S_{pri}( \omega_{k} )}_{1} \\\vdots \\{S_{pri}( \omega_{k} )}_{M}\end{bmatrix}}.}$

As may be appreciated from the foregoing, the above-described estimatoruses multiple measurements (i.e., M noise signals measured by each ofthe primary and auxiliary coils) to estimate the value of theprimary-to-auxiliary transform for multiple frequency bins. This resultsin significantly improved estimates of the PA transform as compared totechniques which rely on a single measurement (i.e., a single signalmeasured by each of the primary and auxiliary coils) to estimate thetransform. Such single-measurement techniques may include scaling andtime-shifting the reference signal before subtraction, which wouldcorrect for a difference in phase between the noise signal as receivedat a primary coil and an auxiliary coil, but (unlike the multiplemeasurement technique described herein) would not correct forfrequency-dependent phase differences.

Another single-measurement technique may include scaling and phaseadjusting the auxiliary noise signal in the frequency domain beforesubtracting it from the signal received at the primary coil. This couldbe accomplished by using the discrete Fourier transform (DFT) of thesignals received by a primary coil and an auxiliary coil. The optimalscaling and phase shift can be determined by a least-squares fit acrossmultiple frequency bins. For example, if S_(pri)(ω_(k)) is the DFT ofthe signal measured on the primary receive coil and S_(aux)(ω_(k)) isthe DFT of the signal measured on an auxiliary coil at the same time, anaverage scaling and phase shift SPF for a subset of frequency bins (inthe range of [k1,k2]) may be computed according to:

${SPF} = {\frac{\sum\limits_{k\; 1}^{k\; 2}\;{{S_{aux}( \omega_{k} )}{S_{pri}( \omega_{k} )}}}{\sum\limits_{k\; 1}^{k\; 2}\;{{S_{aux}( \omega_{k} )}{S_{aux}( \omega_{k} )}}}.}$

Although this single-measurement technique may be used to create afrequency-dependent correction, the method requires a tradeoff betweenfrequency resolution of the correction and accuracy of the estimation ofthe scaling and phase offset. In particular, this “averaging acrossfrequency bins of a single measurement” technique results in poor (e.g.,high-variance, biased) estimation of a PA transform. In contrast, theabove-described multiple measurement technique provides for an unbiasedand low-variance estimator.

As described above, the inventors have appreciated that the use ofmultiple coils may facilitate improved MRI in a number of ways,including more robust noise detection and/or cancellation, acceleratedimage acquisition, etc. In embodiments where multiple primary receivecoils and/or multiple auxiliary sensors are used, all of the sensors maybe the same type or may be of different types. For example, incircumstances where one or more RF coils are used as sensors, none,some, or all of the coils may be shielded. As another example, the coilscan have different sensitivities. When other types of sensors are used,at least some of the characteristics of the sensors and the primaryreceive coil(s) may necessarily be different, though some may be similaror the same.

In some embodiments, multiple auxiliary RF coils and/or primary RF coilsmay be used to accelerate imaging. For example, multiple RF coils usedto sense noise from the same or different noise sources may also be usedto perform parallel MR. In this manner, multiple RF coils may provideboth noise characterization functions as well as accelerated imageacquisition via their use as parallel receive coils.

In some embodiments, as described above, multiple sensors may be used toperform noise compensation in the presence of multiple noise sources. Inan environment having N correlated noise sources, where N is an integergreater than one, the Fourier transforms C_(pri)(ω) and C_(aux)(ω) ofnoise signals c_(pri)(t) and c_(aux)(t), received by a primary coil andan auxiliary sensor can be expressed as:C _(pri)(ω)=H _(pri,1)(ω)C ₁(ω)+H _(pri,2)(ω)C ₂(ω)+ . . . +H_(pri,N)(ω)C _(N)(ω)C _(aux)(ω)=H _(aux,1)(ω)C ₁(ω)+H _(aux,2)(ω)C ₂(ω)+ . . . +H_(aux,N)(ω)C _(N)(ω),where C_(j)(ω); 1≤j≤N, is a Fourier transform of a noise signal from thej^(th) noise source, H_(pri,j)(ω) is a transform between the primarycoil and the j^(th) noise source, and H_(aux,j)(ω) is a transformbetween the auxiliary sensor and the j^(th) noise source. When the ratioH_(pri,j)(ω)/H_(aux,j)(ω) is different for one or more noise sources, itmay not be possible to perform high quality noise compensation by usingonly a single auxiliary sensor. However, multiple auxiliary sensors maybe used to perform noise compensation in this circumstance as describedbelow.

Described below is a non-limiting example of how multiple auxiliarysensors may be used to perform noise compensation for multiple differentnoise sources. Without loss of generality, suppose a MR system has aprimary coil and P auxiliary sensors (where P is any integer greaterthan or equal to 1). Further, suppose that the MR system is deployed inan environment in which there are N different noise sources (where N isan integer greater than or equal to 1). Let H_(ij)(ω) denote thetransform between the i^(th) auxiliary sensor (where 1≤i≤P) and the jthnoise source (where 1≤j≤N). The following set of equations relate theFourier transforms of the signals received by the auxiliary sensors tothe Fourier transforms of the noise signals produced by the noisesources:

${{\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}\begin{bmatrix}C_{1} \\\vdots \\C_{N}\end{bmatrix}} = \begin{bmatrix}C_{{aux},1} \\\vdots \\C_{{aux},P}\end{bmatrix}},$where C_(aux,i); 1≤i≤P, is a Fourier transform of the signal received atthe ith auxiliary sensor, C_(j)(ω); 1≤j≤N is a Fourier transform of anoise signal from the j^(th) noise source, and where the dependence ofall the terms on frequency is not shown explicitly (the (ω) issuppressed for brevity), though it should be appreciated that all theterms in the above matrix equation are functions of frequency.

When the number of auxiliary sensors is greater than or equal to thenumber of noise sources (i.e., P>=N), the above matrix equation may besolved for the noise signals according to:

$\begin{bmatrix}C_{1} \\\vdots \\C_{N}\end{bmatrix} = {{{\{ {\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}^{T}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}} \}^{- 1}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}}^{T}\begin{bmatrix}C_{{aux},1} \\\vdots \\C_{{aux},P}\end{bmatrix}}.}$

If such a solution exists, the correlated noise measured on the primaryreceive coil may be expressed in relation to the measurements obtainedby all of the auxiliary sensors according to:

$C_{pri} = {\begin{bmatrix}H_{{pri},1} & \ldots & H_{{pri},N}\end{bmatrix}{{\{ {\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}^{T}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}} \}^{- 1}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}}^{T}\begin{bmatrix}C_{{aux},1} \\\vdots \\C_{{aux},P}\end{bmatrix}}}$

A multi-channel transform H_(MPA) may be defined according to:

$H_{MPA} = {\begin{bmatrix}H_{{PA},1} & \ldots & H_{{PA},P}\end{bmatrix} = {\begin{bmatrix}H_{{pri},1} & \ldots & H_{{pri},N}\end{bmatrix}{{\{ {\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}^{T}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}} \}^{- 1}\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}}^{T}.}}}$

It may then be seen that the noise measured by the primary receive coilis a linear combination of the noise signals measured on all theauxiliary coils:

$\begin{matrix}{C_{pri} = {{\begin{bmatrix}H_{{PA},1} & \ldots & H_{{PA},P}\end{bmatrix}\begin{bmatrix}C_{{aux},1} \\\vdots \\C_{{aux},P}\end{bmatrix}}.}} & (3)\end{matrix}$

Thus, given noise signals measured by P auxiliary sensors (e.g., theFourier transforms of which are given by C_(aux,i) for 1≤i≤P), the aboveequation may be used to estimate the noise signal received at theprimary receive coil (e.g., the Fourier transform of which is given byC_(pri)). In turn, the estimated noise signal may be subtracted from theoverall signal measured by the primary receive coil (which signal wouldhave both an MR signal component and a noise component) to perform noisesuppression.

However, to use the above equation (3), an estimate of the multichannelprimary-to-auxiliary transform H_(MPA)=[H_(PARC,1) . . . H_(PARC,P)] isneeded. This may be achieved in any suitable way and, in someembodiments, may be done by making multiple measurements using theprimary receive coil and the auxiliary sensors (e.g., at a time whenthere is no MR signal present) and using these measurements to estimatethe multichannel primary-to-auxiliary transform. For example, given Mmeasurements of noise signals at each of the P auxiliary sensors and theprimary receive coil, the H_(MPA) may be estimated for each frequencycomponent ω_(k) (where k is an index over frequency bins) using leastsquares estimation according to:

${\begin{bmatrix}{H_{{PA},1}( \omega_{k} )} \\\vdots \\{H_{{PA},P}( \omega_{k} )}\end{bmatrix} = {\{ {\begin{bmatrix}{S_{{aux},1}( \omega_{k} )}_{1} & \ldots & {S_{{aux},P}( \omega_{k} )}_{1} \\\vdots & \ddots & \vdots \\{S_{{aux},1}( \omega_{k} )}_{M} & \ldots & {S_{{aux},P}( \omega_{k} )}_{M}\end{bmatrix}^{T}\begin{bmatrix}{S_{{aux},1}( \omega_{k} )}_{1} & \ldots & {S_{{aux},P}( \omega_{k} )}_{1} \\\vdots & \ddots & \vdots \\{S_{{aux},1}( \omega_{k} )}_{M} & \ldots & {S_{{aux},P}( \omega_{k} )}_{M}\end{bmatrix}} \}^{- 1} \times \begin{bmatrix}{S_{{aux},1}( \omega_{k} )}_{1} & \ldots & {S_{{aux},P}( \omega_{k} )}_{1} \\\vdots & \ddots & \vdots \\{S_{{aux},1}( \omega_{k} )}_{M} & \ldots & {S_{{aux},P}( \omega_{k} )}_{M}\end{bmatrix}^{T} \times \begin{bmatrix}{S_{pri}( \omega_{k} )}_{1} \\\vdots \\{S_{pri}( \omega_{k} )}_{M}\end{bmatrix}}},$where S_(aux,i)(ω_(k))_(m) represents the value of the kth frequency binof the Fourier transform of the mth measured signal obtained by the ithauxiliary sensor, and where S_(pri)(ω_(k))_(m) represents the value ofthe kth frequency bin of the Fourier transform of the mth measuredsignal obtained by the primary receive coil. This least-squares approachprovides the most complete correction when the columns of the followingmatrix are as orthogonal as possible to one another:

$\begin{bmatrix}H_{11} & \ldots & H_{1N} \\\vdots & \ddots & \vdots \\H_{P\; 1} & \ldots & H_{PN}\end{bmatrix}.$

Put another way, each auxiliary sensor may detect some or all of thedifferent noise sources in a unique way from other auxiliary sensors. Inorder to correct for the presence of near field sources, multiplesensors may be placed in different locations to be more or lesssensitive to some of the noise sources. In some embodiments, multiplesensors may be oriented orthogonally to one another (e.g., one sensormay be oriented in an “X” direction, another sensor may be oriented inthe “Y” direction, and another sensor may be oriented in a “Z”direction). In this way, each vector of the time varying interferencefields may be captured. It may also be beneficial to use one or moreantennas as an auxiliary sensor to provide another orthogonalmeasurement.

It should be appreciated that the techniques described herein facilitatedetecting noise in the environment of an MRI system using any numberand/or type of sensor suitable for detecting noise produced byrespective noise sources. As a result, noise from a variety of sourcesthat may impact the performance of the MRI system may be detected andused to suppress and/or eliminate noise from MR signals detected by theMRI system during operation. Because techniques described herein operateon the particular noise environment of the MRI system, a noise reductionsystem employing noise suppression techniques described hereinfacilitate deployment of an MRI system wherever the system may beneeded, eliminating the requirement that the system be installed inspecially shielded rooms. The ability to dynamically adapt to changingnoise environments facilitates development of MRI systems that can bedeployed in generally noisy environments, including environments wherenoise sources may change over time. Because techniques described hereincan be utilized during operation of the MRI system, the noiseenvironment can be characterized dynamically so that it reflects thesame noise environment to which the system is currently being exposed.These noise suppression and/or avoidance techniques permit the MRIsystem to operate in almost any environment and to dynamically adapt toand compensate for electromagnetic noise present, enabling a portableMRI system that can be transported to wherever the patient is located toperform the needed diagnostic, surgical or monitoring procedure.

A noise reduction system may include additional techniques to increasethe SNR of a portable MRI system by reducing system noise, for example,by reducing inductive coupling between adjacent or neighboring RF coilsin a multi-coil transmit/receive system. According to some embodiments,multiple coils can be used to both improve SNR and to facilitate noisesuppression. For example, a collection of RF coils, which may be eitherRF signal coils (e.g., primary RF coils), RF noise coils (e.g.,auxiliary RF coils) or both, may be arranged at different locations andorientations to detect a comprehensive RF field that can becharacterized and compensated for using any of the noise suppressiontechniques discussed herein. According to some embodiments, a portableMRI system comprises multiple transmit/receive coils to improve the SNRof image acquisition. For example, a portable MRI system may comprise 2,4, 8, 16, 32 or more RF receive coils to improve the SNR of MR signaldetection.

In general, RF coils are tuned to increase coil sensitivity at afrequency of interest. However, inductive coupling between adjacent orneighboring coils (e.g., RF coils sufficiently proximate one another)degrades the sensitivity of tuned coils and significantly reduces theeffectiveness of the collection of RF coils. Techniques forgeometrically decoupling neighboring coils exist but place strictconstraints on coil orientation and position in space, reducing theability of the collection of RF coils to accurately detect the RF fieldand, as a consequence, degrading the noise rejection performance. Toaddress the negative impact of inductive coupling between coils, theinventors have utilized coil decoupling techniques that reduce theinductive coupling between radio frequency coils in multi-coiltransmit/receive systems. For example, FIGS. 43A and 43B illustratepassive decoupling circuits configured to reduce inductive couplingbetween radio frequency coils in a multi-coil transmit/receive system,in accordance with some embodiments. Passive decoupling circuit 4300 amay be configured to decouple RF noise coils, for example, RF noisecoils positioned outside the field of view of the MRI system that arenot subjected to the relatively intense transmit B₁ field produced bythe RF transmit system (i.e., one or more RF transmit coils). In thiscontext, inductor L1 represents an RF coil configured to detectelectromagnetic noise in the environment that is tuned by the circuitformed by capacitors C1, C2 and C3. Capacitors and inductors arearranged to provide a balanced differential circuit to reduce commonmode noise. The tank circuit formed by L2, L3, C3 and C4 is configuredto have a high impedance to ensure that the current through L1 remainssmall. Appropriate selection of the values for the L-C network ensuresthat the current passing through L1, while remaining small, hassufficient SNR for measurement at the differential output (Vout−, Vout+)of the LNA to characterize electromagnetic noise in the environment withadequate sensitivity. Equivalent impedance at the LNA input is given by:

$Z_{eq} = {( \frac{C_{4}}{C_{3}} )^{2}R}$

In the above equation, R is the equivalent losses of the primaryinductance L1. Capacitor and inductor values can be chosen to attainoptimal noise impedance of the LNA used for detection. FIG. 43Billustrates a passive decoupling circuit 4300 b configured to decoupleRF coils that may be subjected to B₁ transmit fields. In particular, L1may represent an RF signal coil within the field of view of the MRIsystem. Passive decoupling network 4300 b may be similar to passivedecoupling network 4300 a in some respects, but differs in that diodeD1, capacitor C3 and inductors L2 and L3 operate as a transmit/receiveswitch that isolates the RF coil (represented as inductor L1) from theLNA when RF signals are being transmitted by one or more RF transmitcoils. Specifically, the L-C network is divided into two networkportions by the transmit/receive switch to protect sensitive electronicsduring RF transmit cycles. During a transmit pulse, diode D1 is turnedon to create a short circuit, isolating the RF signal coil from thereceive electronics. The resulting L-C network provides a tank circuitwith a high impedance that ensures that the current in L1 remains small.During receive cycles, diode D1 is turned off and the RF coil isconnected to the LNA and tuned by the resulting balanced tank circuitconfigured to limit the current through L1, while allowing forsufficient signal to be detected at the differential outputs of the LNA.Thus, the RF coil is connected to a first tank circuit during transmitcycles and a second tank circuit during receive cycles of a pulsesequence. Equivalent impedance at the LNA input is then:

$Z_{eq} = {( \frac{L_{4}}{L_{2}} )^{2}R}$

Conventional decoupling circuits often use PIN diodes to isolate thereceive electronics from the RF signal coil. However, PIN diodessuitable for performing this function in a decoupling circuit requireapproximately 1 A of current to turn the diode on. As an example, atransmit/receive coil system having eight receive coils may require onthe order of 8 A of current to decouple the receive coils from the RFsignal coil(s) for each transmit and receive cycle of an imageacquisition pulse sequence. Accordingly, over the span of an imageacquisition protocol, substantial power is consumed by the decouplingcircuits of the RF transmit/receive system. The inventors recognizedthat PIN diodes could be replaced by Gallium Nitride (GaN) field effecttransistors (FETs) to reduce the power consumption of the RFtransmit/receive system. In particular, GaN FETs require on the order ofmilliamps to turn on, reducing the power consumption by several ordersof magnitude. In addition, the capacitance of the GaN FETs when turnedon are small compared to PIN diodes, reducing negative impact on thebalanced circuit. According to some embodiments, diode D1 in decouplingcircuit 4300 b is replaced with one or more GaN FETs, thereby reducingthe power consumption of the RF transmit/receive system.

FIG. 43C illustrates a circuit 4300 for a RF receive coil 4326 using GaNFETs 4335 to couple and decouple the receive electronics from the RFcoil, in accordance with some embodiments. In FIG. 43C, a receive coil4326 is connected to a resonant circuit 4333 and to receive circuitry4340 (e.g., preamplifiers such as linear amplifiers) configured toreceive and deliver signals detected by receive coil 4326. Duringtransmit cycles (e.g., during transmission of RF pulses by one or moreRF transmit coils), receive coil 4326 is detuned to protect receivecircuitry 4340, which could be damaged if RF transmit signals from RFtransmit coil(s) were to couple to receive coil 4326 and propagate toreceive circuitry 4340. As discussed above, conventional circuits oftenemploy PIN diodes to detune or decouple the receive coil from thereceive circuitry. Circuit 4300 includes decoupling circuitry that usesGaN FETs 4335 to detune receive coil 4326 so as to decouple the receivecoil from receive circuitry 4340. In particular, during transmit cycles,GaN FETs 4326 are turned on (i.e., closed to create a short circuitbetween terminals) to switch inductor 4337 into the circuit to detuneresonant circuit 4333 so that RF transmit pulses do not couple toreceive coil 4326. During receive cycles, GaN FETs 4335 are turned off(i.e., opened to create an open circuit between terminals) to removeinductor 4337 from resonant circuit 4333 so that receive coil willcouple to MR signals emitted in response to the RF transmit pulses. Asdiscussed above, GaN FETs require substantially less power to turn oncompared to conventional diodes such as PIN diodes, conserving power oneach transmit/receive cycle (e.g., reducing the power consumption fromapproximately 1 A to milliamps for each receive coil in the RFtransmit/receive system).

FIG. 43D illustrates an active decoupling circuit configured to reduceinductive coupling between radio frequency coils in a multi-coiltransmit/receive system, in accordance with some embodiments. In thedecoupling circuit illustrated in FIG. 43D, inductor L1 represents an RFcoil configured to measure an NMR signal. The RF coil is tuned viacapacitor C3 connected in parallel to L1, and the differential outputsVout−, Vout+ of the LNA measure the NMR signal sensed by the RF coil.The differential output of the LNA are also fed back to a secondinductor L2 via resistors R1 and R2. The feedback circuit causes currentflowing through inductor L2 to couple negative flux into L1 in responseto signals, thus reducing the current flowing through L1 andconsequently mitigating the inductive coupling effect on other nearby RFcoils. L2 may be provided at a desired distance from L1 and the resistorvalues of R1 and R2 can be chosen so that the current through L2achieves the desired current reduction in L1. Decoupling circuit 4300 creduces the number of circuit elements required, thereby reducing thecost and complexity of the decoupling circuit.

The use of decoupling circuits, such as the decoupling circuitsillustrated in FIGS. 43A, 43B, 43C and 43D facilitates increasing SNRand mitigates the impact of inductive coupling on the noise rejectionperformance of a noise reduction system in a multi-coil transmit/receivesystem. In addition, the decoupling circuit illustrated in FIG. 43Cprovides a low power transmit/receive switch that reduces the powerconsumption of decoupling and coupling the RF receive coils duringtransmit and receive cycles, respectively, via the use of GaN FETs(e.g., instead of PIN diodes and the like). Accordingly, the RF coilsystem may be operated with reduced power consumption. It should beappreciated that other decoupling circuits may be used, as the aspectsare not limited in this respect.

According to some embodiments, noise from various sources arecharacterized using a combination of the above described techniques todetermine a multi-channel transform that can be used to suppress oreliminate noise from the various noise sources. Noise measurements maybe obtained during operation of the MRI system so that a multi-channeltransform may be determined dynamically, allowing for noise suppressionthat adapts to the changing noise environment of the MRI system.However, noise in the environment may be characterized upon systemstart-up, when the system is moved to a different location and/or uponthe occurrence of any event, and the characterized noise used tosuppress and/or eliminate noise in acquired MR signals, as thetechniques described herein can be applied as desired. Any other noisesuppression techniques may also be utilized to facilitate operation ofan MRI system outside a specially shielded room, tent or enclosureand/or where shielding of the imaging region is otherwise limited orabsent, thus allowing for portable MRI.

It should be appreciated that the these noise suppression techniquesfacilitate detecting noise in the environment of an MRI system using anynumber and/or type of sensor suitable for detecting noise produced byrespective noise sources. As a result, noise from a variety of sourcesthat may impact the performance of the MRI system may be detected andused to suppress and/or eliminate noise from MR signals detected by theMRI system during operation. Because these techniques operate on theparticular noise environment of the MRI system, a noise reduction systememploying these noise suppression techniques facilitate deployment of anMRI system wherever the system may be needed, eliminating therequirement that the system be installed in specially shielded rooms.The ability to dynamically adapt to changing noise environmentsfacilitates development of MRI systems that can be deployed in generallynoisy environments, including environments where noise sources maychange over time. Because the described noise suppression techniques canbe utilized during operation of the MRI system, the noise environmentcan be characterized dynamically so that it reflects the same noiseenvironment to which the system is currently being exposed. These noisesuppression and/or avoidance techniques permit the MRI system to operatein almost any environment and to dynamically adapt to and compensate forelectromagnetic noise present, enabling a portable MRI system that canbe transported to wherever the patient is located to perform the neededdiagnostic, surgical or monitoring procedure.

It should be further appreciated that a noise reduction system mayinclude any one or more noise suppression, rejection and/or avoidancetechniques described herein (e.g., one or more of dynamic noisesuppression, rejection and/or avoidance techniques, one or moredecoupling circuits to reduce inductive coupling, etc.) to facilitateoperation of the portable MRI system in virtually any room and withvirtually any device-level shielding configuration. As discussed above,conventional MRI systems operate in specially shielded rooms thatprovide an encompassing shielded space. As a result, MRI systemsoperating in specially shielded rooms have shielding for substantially100% of the imaging region. MRI systems that operate within moveabletents or cages also have comprehensive shielding of the imaging regionthat endeavor to provide as close to 100% shielding of the imagingregion as is practicable. To achieve portability, MRI systems accordingto some embodiments are configured to operate outside specially shieldedrooms, tents or cages with varying levels of device-level shielding(e.g., shielding some fraction of the imaging region), including no, orsubstantially no, shielding of the imaging region.

The amount of electromagnetic shielding for an imaging region can beviewed as a percentage of the maximum solid angle, subtending theimaging region from its center, for which shielding is provided.Specifically, providing shielding for 100% of an imaging region meansthat electromagnetic shielding for at least the operating spectrum isprovided over the maximum solid angle 4π steradian (sr) about theimaging region. Similarly, providing shielding for less than 75% of theimaging region means that electromagnetic shielding for at least theoperating spectrum provides less than 0.75(4π) sr solid angle coverageof the imaging region, and so on. Accordingly, a specially shielded roomprovides shielding for substantially 100% of the imaging region for anMRI system deployed within the shielded room because shielding isprovided over substantially the maximum solid angle of 4π sr. Similarly,moveable tents or cages are designed to provide shielding for as closeto as 100% of the imaging region as is practicable.

The percentage of electromagnetic shielding of an imaging region of anMRI system refers to the total amount of shielding that protects theimaging region, including electromagnetic shielding provided viaspecially-shielded rooms, tents, cages, etc., as well as device-levelelectromagnetic shielding (e.g., electromagnetic shields coupled to thehousing of the MRI device that provide electromagnetic shielding for theimaging region). Thus, the portable MRI systems illustrated in FIGS. 16,39A-C, 40A and 40B, when operated outside of specially-shielded rooms orcages, have less than 100% shielding of their respective imaging regionsand, in some configurations, have substantially less than 100%shielding. Providing shielding for less than 100% of the imaging regionis referred to herein as providing shielding for a fraction of theimaging region, which fraction may be quantified by a specificpercentage or percentage range. For example, the electromagnetic shieldsillustrated in FIGS. 16, 39A-C, 40A and 40B may be adjusted to provideshielding for different fractions of the imaging region (e.g., varyingdegrees of shielding), such as at least between approximately 85% andapproximately 50% (e.g., at approximately 85% or less, approximately 75%or less, approximately 65% or less, etc.).

It should be understood that providing shielding for a fraction of theimaging region refers to instances in which providing less than 100%shielding is intentional and/or by design (e.g., to provide access to oraccommodate a patient in an MRI system operated outside a speciallyshielded room, tent or cage). In practice, shielding techniques areoften imperfect and therefore may provide less than 100% shielding eventhough the intent is to provide 100% shielding for the imaging region(at least for the operating spectrum). For example, doors that are leftopen or ajar in specially shielded rooms, gaps in tents that gounnoticed, or openings that are not fully closed during imaging, etc.,may result in less than 100% shielding even though the intent is toprovide full coverage. Imperfect shielding material or construction mayalso result in unintentionally having less than 100% shielding.Providing shielding for a fraction of the imaging region should not beinterpreted to cover these situation, as it refers to circumstanceswhere the fractional coverage is intentional and/or by design.

FIGS. 44A-C illustrate a portable MRI system having different amounts ofdevice-level shielding about the imaging region, in accordance with someembodiments. FIG. 44A illustrates a portable MRI system having shields5065 that partially shield the imaging region 5095. For example, shields5065 may be incorporated into slides 5060 that can be configured andpositioned as desired to provide shielding around approximately 50% ofthe opening to imaging region 5095. FIG. 44B illustrates another exampleof a portable MRI system having shield 5165 that provides a lesserdegree of shielding for imaging region 5195. For example, slide 5160 maybe positioned as desired to provide shielding around approximately 25%of the opening to imaging region 5195. FIG. 44C illustrates an exampleof a portable MRI system without shields around the imaging region 5295,providing an alternative having substantially no device-level shieldingfor the open imaging region.

FIG. 44D illustrates a portable MRI system 4400 utilizing a furthertechnique for electromagnetic shielding of the imaging region of thesystem, in accordance with some embodiments. In particular, in theembodiment illustrated in FIG. 44D, shielding from electromagneticinterference is achieved via one or more conductive strips connectingupper and lower portions of the B₀ magnet of the portable MRI system toform a conductive loop that counteracts at least some electromagneticradiation that would otherwise result in interference. In the embodimentillustrated in FIG. 44D, conductive strip 4465 is electrically coupledto upper portion 4400 a, lower portion 4400 b and may also be connectedto ground. In the embodiment illustrated in FIG. 44D, conductive strip4465 is formed by a conductive braid, providing a flexible strip ofmaterial that can be coupled to the B₀ magnet with relative ease andconvenience. However, conductive strip 4465 may be constructed orcomposed of any conductive material in any suitable form, some examplesof which are described in further detail below.

The exemplary portable MRI system 4400 illustrated in FIG. 44D includesa ferromagnetic yoke 4420 that provides a magnetic path between uppermagnet 4410 a and lower magnet 4410 b to capture and direct the magneticfield produced by the respective magnets to increase the magnetic fluxdensity within the imaging region. In particular, similar to theexemplary yokes described in connection with FIGS. 2A-B, 3A and 16, yoke4420 comprises a frame and upper and lower plates formed using asuitable ferromagnetic material or combination of materials (e.g., iron,steel, etc.). The upper and lower plates are coupled to the upper andlower magnets, respectively, to form a “magnetic circuit” that capturesat least some of the magnetic field produced by the magnets and directsthe captured magnetic fields via the “magnetic circuit” to increase theflux density within the imaging region of the MRI device.

The inventors have recognized that coupling conductive strip 4465 to theplates of the yoke forms a conductive loop in which current is inducedby electromagnetic radiation propagating in directions through theconductive loop. This induced current will in turn produce anelectromagnetic field that counteracts at least some of theelectromagnetic radiation that induced the current and/orelectromagnetic radiation similarly propagating through the loop. Inthis manner, electromagnetic interference can be reduced by thecounteracting electromagnetic field produced by current induced in theconductive loop formed by the conductive strip 4465 and yoke 4220.Accordingly, the suppression of electromagnetic interference may beimproved by the addition of further conductive strips forming additionalconductive loops to produce counteracting electromagnetic fields whenambient electromagnetic radiation induces current in the respectiveconductive loop. In particular, as more conductive loops are added atdifferent orientations, the resulting conductive loops will beresponsive to more of the electromagnetic radiation present in theenvironment.

It should be appreciated that any number of conductive strips may beattached or affixed to the B₀ magnet to provide electromagneticshielding. According to some embodiments, one or more additional strips4465 connecting components of the B₀ magnet to ground may be providedabout the imaging region to increase the amount of shielding arranged toprotect the imaging region from electromagnetic interference (e.g., toincrease the percentage of electromagnetic shielding for the imagingregion). For example, a conductive strip shield may be attached every180°, every 90°, every 45°, every 30° or at any other interval, eitherregularly or irregularly spaced about the imaging region, to provide adesired degree of electromagnetic shielding. It should be appreciatedthat any number of conductive strips may be used to achieve a desiredpercentage of shielding and/or to deliver a desired compromise betweenopenness of the imaging region and comprehensiveness of the shieldingfor the imaging region, as discussed in further detail below.

While the conductive strip 4465 illustrated in FIG. 44D is made from aflexible material, one or more conductive strips may be formed in otherways, for example, as a rigid conductive strip, bar, rod or handle (orother suitable geometry) that electrically connects the magnets formingthe B₀ magnet of the MRI system to ground. In this respect, one or moreconductive strips may be arranged to serve as a handle to assist inmoving the portable MRI system, to facilitate rotating the device or toassist in tilting the B₀ magnet (e.g., in conjunction with a goniometricmember, examples of which are described in connection with FIGS. 45-47below) in addition to providing electromagnetic shielding. It should beappreciated that different types of conductive strips may be used incombination (e.g., one or more flexible strips and/or one or more rigidstrips) to provide electromagnetic shielding for the MRI system, as theaspects are not limited in this respect.

According to some embodiments, one or more conductive strips areconfigured to be removable so that conductive strips can be added andremoved as desired, facilitating configurable strip shielding thatprovides a flexible approach to accommodate different operatingenvironments, different imaging circumstances and/or different levels ofclaustrophobic affliction or unease of the patient. To facilitateconfigurable shielding in this respect, the housing for the magnets mayinclude a plurality of fastening mechanisms (e.g., snaps, connectors,inserts or other mechanisms) that allow for removable attachment ofconductive strips to the housing and that electrically couple themagnets to the conductive strips and to ground when a conductive stripshield is connected to the housing via a respective fastening mechanism.Fastening mechanisms may be arranged at any desired location and at anynumber of locations to provide flexibility in where and how manyconductive strips may be attached to the device. Additionally, thefastening mechanisms themselves may be made to be moveable so that oneor more conductive strips coupled to the system via the fasteningmechanisms may be adjusted (e.g., rotated about the imaging region). Inthis manner, conductive strips may be added, removed and/or theirpositions adjusted as needed to provide a desired shieldingconfiguration in a desired amount (e.g., to provide shielding for adesired percentage of the imaging region).

Providing a plurality of fastening mechanisms that allow removablestrips to be attached and removed at a number of locations about theimaging region allows the imaging region to remain essentially openwhile positioning a patient within the imaging region. After the patienthas been positioned within the imaging region, a desired number ofconductive strips may be attached to the B₀ magnet via the plurality offastening mechanisms to achieve a desired degree of shielding, toaddress the electromagnetic environment in which the MRI system isoperating, to facilitate a particular imaging protocol and/or toaccommodate a patient who may be susceptible to claustrophobia (e.g.,conductive strips may be added only while the patient remainscomfortable with the openness of the MRI system). Accordingly, stripshielding techniques may provide a flexible, configurable approach toelectromagnetic shielding, facilitating the ability to deploy portableMRI systems in a variety of environments and for a variety ofapplications and circumstances.

There are a number of benefits to reducing the shielding provided aroundthe imaging region (e.g., using any of the shielding techniquesdescribed herein), including a reduction in cost and complexity of thesystem and improvements in accessibility to the imaging region both withrespect to positioning a patient for imaging, as well as increasedaccessibility for medical personnel who may need to perform other tasksrequiring access to the patient while the patient remains positionedwithin the system. In addition, reducing the shielding around theimaging region maximizes the openness of the MRI system to improve theexperience of patients who are susceptible to feelings ofclaustrophobia. In this manner, the applicability of portable MRI may befurther increased from a cost and/or flexibility perspective.

According to some embodiments, device-level shields are removable suchthat the amount of shielding provided may be selected in view of theparticular circumstances, such as the required accessibility to thepatient and/or imaging region for a given procedure, the severity of apatient's claustrophobia, the particular noise environment, etc. Forexample, slides carrying shields may be configured to be attached andremoved from the B₀ magnet, allowing for a portable MRI device to beselectively and dynamically configured as desired (e.g., to allow aportable MRI system to be configured with the amount of shielding andaccessibility illustrated in FIGS. 40 and 44A-C, which have three, two,one and zero slides/shields attached, respectively). In this manner, aportable MRI device can take advantage of the shielding andaccessibility aspects of the different possible configurations, allowingthe portable MRI to be optimized in this respect for given proceduresand/or particular patients. According to other embodiments, the numberof shields or the amount of shielding for a given portable MRI systemmay be fixed, which may allow for reductions in cost and complexity, butmay also decrease the flexibility of the system from ashielding/accessibility perspective.

As discussed above, the inventors have developed noise reduction systemsthat allow a portable MRI device to operate in different noiseenvironments (e.g., in unshielded or partially shielded rooms) and tooperate with varying amounts of device-level shielding. A portable MRIsystem may include a noise reduction system that includes any one orcombination of the noise suppression, avoidance and/or reductiontechniques described herein, as the aspects are not limited in thisrespect. For example, a noise reduction system may employ one or more ofthe noise suppression and/or avoidance techniques described herein,allowing for dynamic noise suppression and/or avoidance that compensatesfor a given noise environment and/or that works in concert with thevariable amounts of device-level shielding provided by portable MRIsystems having configurable shields (e.g., the portable MRI systemsillustrated in FIGS. 40 and 44A-C), including no or substantially nodevice-level shielding about the imaging region. A noise reductionsystem may also include coil decoupling networks to reduce the noiseresulting from inductive coupling between radio frequency coils inmulti-coil transmit/receive systems at any level of shielding provided.It should be appreciated that a noise reduction system may include anyone or combination of techniques described herein, as the aspects arenot limited in this respect.

As shown in FIGS. 16, 39A-C, 40A and 40B, the portable MRI is configuredso that the B₀ magnet can be tilted at a desired incline. In manyinstances, a patient may not be able to lie flat, for example, due torisks associated with increased hydrostatic pressure in the brain. Theinventors have developed a portable MRI device having a positioningmechanism that allows the B₀ magnet to be rotated, for example, aboutits center of mass. Thus, if a patient or a particular portion of apatient's anatomy needs to be supported at an incline, the positioningmechanism can be engaged to rotate or tilt the B₀ magnet to achieve thedesired incline. According to some embodiments, the positioningmechanism can be manually engaged to rotate or tilt the B₀ magnet byhand, facilitating quick and easy configuring of the MRI system at thedesired incline.

FIGS. 45A-45D illustrate different views of a positioning mechanism thatemploys a positioning goniometer or goniometric stage 4590 that allowsthe magnet to be rotated about a fixed axis (e.g., the axis through ornear the center of mass of the B₀ magnet). As illustrated in FIG. 45A, agoniometric stage 4590 is rotatably coupled to the bottom of a lowerportion of the B₀ magnet to allow the B₀ magnet to be rotated about itscenter of mass 4591, as shown by direction arrows 4593 in FIG. 45A.Goniometric stage 4590 includes a number of holes or bores 4595configured to accommodate a locking member (e.g., a locking pin) thatlocks the mechanism in place at a desired angle, as discussed in furtherdetail below. Rotating the B₀ magnet via goniometric stage 4590 effectsa tilt that provides an inclined supporting surface for the patientanatomy being imaged, as illustrated in FIG. 40A.

FIG. 45B illustrates a side view of the B₀ magnet and the goniometricstage 4590. Goniometric stage 4590 includes a release mechanism 4594that engages and disengages a locking pin 4596 from holes provided on afixed or stationary member of the goniometric stage 4590 (e.g., holes4595 illustrated in FIG. 45A). To rotate the B₀ magnet, releasemechanism 4594 is pressed in an upward direction to disengage thelocking pin 4596 from the hole in which it is currently positioned. Forexample, handle 4592 allows a user to place a hand on the handle andsqueeze release mechanism 4594 towards the handle to release the lockingpin 4596, as discussed in further detail below in connection with FIGS.46A and 46B. With the locking pin 4596 disengaged, the B₀ magnet maythen be rotated or tilted to the desired incline using handle 4592. Oncethe B₀ magnet has been rotated to the desired angle, release mechanism4594 may be released so that the locking pin 4596 engages with acorresponding hole at the new position, locking the mechanism in placeat the desired angle. FIGS. 45C and 45D illustrate the goniometric stage4590 coupled to the bottom side of the lower magnet apparatus 4510 andbase 4550.

FIGS. 46A and 46B illustrate a closer view of the exemplary goniometricstage 4590 discussed in the foregoing. As shown, release mechanism 4594is rotatably coupled to moveable stage component 4590 a via axle 4599.When force is applied to release mechanism 4594 in the direction shownby arrow 4597 (e.g., by gripping handle 4592 and release mechanism 4594with a hand and squeezing or lifting the release mechanism towards thehandle), release mechanism 4594 rotates about axle 4599 and raisesportion 4598 to lift locking pin 4596 out of hole 4595 c in which it iscurrently positioned. When locking pin 4596 is lifted from the hole asshown by the phantom lines in FIG. 46B, moveable stage component 4590 ais released from its locked position and allowed it to slide withinstationary stage component 4590 b. When the moveable stage component4590 a is moved to its desired location, release mechanism 4594 can berelease to lock moveable stage component 4590 a into the desiredposition. For example, a spring mechanism may be coupled to the lockingpin so that when release mechanism 4594 is released, the spring forcecauses the locking pin 4596 to return to its locked position. Whileexemplary goniometric stage 4590 includes four holes (e.g., holes 4595a, 4595 b, 4595 c and 4595 d), any number of holes at any location maybe provided to provide a desired granularity to the angles at which theB₀ magnet can be positioned, as the aspects are not limited in thisrespect. It should be appreciated that goniometric stage 4590 allows theB₀ magnet to be rotated without movement of the center of mass,permitting the magnet to be rotated by hand. However, other mechanismsthat rotate the center of mass may also be used, as the aspects are notlimited in this respect.

FIG. 47 illustrates the results of a 3 minute brain scan using aportable MRI system incorporating aspects of the techniques describedherein (e.g., low field MRI system 1900, 3900, 4000, etc.) operatingwith a B₀ magnetic field having a field strength of approximately 50 mT.The proton density images were obtained using a balanced steady statefree precession (bSSFP) pulse sequence and have 2.4×2.2×5 mm resolution.FIG. 48 illustrates the results of a 14 minute brain scan from aportable MRI system operating at a field strength of approximately 50 mTusing a bSSFP pulse sequence. The resolution of the proton densityimages in FIG. 47 is 1.7×1.7×4 mm. FIG. 49 illustrates the result of a15 minute brain scan from a portable MRI system operating at a fieldstrength of approximately 50 mT using a T2 fluid-attenuated inversionrecovery (FLAIR) pulse sequence. The resolution of the T2 images in FIG.49 are 2×2×5 mm. FIG. 50 illustrates a 15 minute scan of the knee usinga portable MRI system operating at approximately 50 mT using a bSSFPpulse sequence. The resolution of the proton density images in FIG. 50is 1.7×1.7×3 mm.

Having thus described several aspects and embodiments of the technologyset forth in the disclosure, it is to be appreciated that variousalterations, modifications, and improvements will readily occur to thoseskilled in the art. Such alterations, modifications, and improvementsare intended to be within the spirit and scope of the technologydescribed herein. For example, those of ordinary skill in the art willreadily envision a variety of other means and/or structures forperforming the function and/or obtaining the results and/or one or moreof the advantages described herein, and each of such variations and/ormodifications is deemed to be within the scope of the embodimentsdescribed herein. Those skilled in the art will recognize, or be able toascertain using no more than routine experimentation, many equivalentsto the specific embodiments described herein. It is, therefore, to beunderstood that the foregoing embodiments are presented by way ofexample only and that, within the scope of the appended claims andequivalents thereto, inventive embodiments may be practiced otherwisethan as specifically described. In addition, any combination of two ormore features, systems, articles, materials, kits, and/or methodsdescribed herein, if such features, systems, articles, materials, kits,and/or methods are not mutually inconsistent, is included within thescope of the present disclosure.

The above-described embodiments can be implemented in any of numerousways. One or more aspects and embodiments of the present disclosureinvolving the performance of processes or methods may utilize programinstructions executable by a device (e.g., a computer, a processor, orother device) to perform, or control performance of, the processes ormethods. In this respect, various inventive concepts may be embodied asa computer readable storage medium (or multiple computer readablestorage media) (e.g., a computer memory, one or more floppy discs,compact discs, optical discs, magnetic tapes, flash memories, circuitconfigurations in Field Programmable Gate Arrays or other semiconductordevices, or other tangible computer storage medium) encoded with one ormore programs that, when executed on one or more computers or otherprocessors, perform methods that implement one or more of the variousembodiments described above. The computer readable medium or media canbe transportable, such that the program or programs stored thereon canbe loaded onto one or more different computers or other processors toimplement various ones of the aspects described above. In someembodiments, computer readable media may be non-transitory media.

The terms “program” or “software” are used herein in a generic sense torefer to any type of computer code or set of computer-executableinstructions that can be employed to program a computer or otherprocessor to implement various aspects as described above. Additionally,it should be appreciated that according to one aspect, one or morecomputer programs that when executed perform methods of the presentdisclosure need not reside on a single computer or processor, but may bedistributed in a modular fashion among a number of different computersor processors to implement various aspects of the present disclosure.

Computer-executable instructions may be in many forms, such as programmodules, executed by one or more computers or other devices. Generally,program modules include routines, programs, objects, components, datastructures, etc. that perform particular tasks or implement particularabstract data types. Typically the functionality of the program modulesmay be combined or distributed as desired in various embodiments.

Also, data structures may be stored in computer-readable media in anysuitable form. For simplicity of illustration, data structures may beshown to have fields that are related through location in the datastructure. Such relationships may likewise be achieved by assigningstorage for the fields with locations in a computer-readable medium thatconvey relationship between the fields. However, any suitable mechanismmay be used to establish a relationship between information in fields ofa data structure, including through the use of pointers, tags or othermechanisms that establish relationship between data elements.

The above-described embodiments of the present invention can beimplemented in any of numerous ways. For example, the embodiments may beimplemented using hardware, software or a combination thereof. Whenimplemented in software, the software code can be executed on anysuitable processor or collection of processors, whether provided in asingle computer or distributed among multiple computers. It should beappreciated that any component or collection of components that performthe functions described above can be generically considered as acontroller that controls the above-discussed function. A controller canbe implemented in numerous ways, such as with dedicated hardware, orwith general purpose hardware (e.g., one or more processor) that isprogrammed using microcode or software to perform the functions recitedabove, and may be implemented in a combination of ways when thecontroller corresponds to multiple components of a system.

Further, it should be appreciated that a computer may be embodied in anyof a number of forms, such as a rack-mounted computer, a desktopcomputer, a laptop computer, or a tablet computer, as non-limitingexamples. Additionally, a computer may be embedded in a device notgenerally regarded as a computer but with suitable processingcapabilities, including a Personal Digital Assistant (PDA), a smartphoneor any other suitable portable or fixed electronic device.

Also, a computer may have one or more input and output devices. Thesedevices can be used, among other things, to present a user interface.Examples of output devices that can be used to provide a user interfaceinclude printers or display screens for visual presentation of outputand speakers or other sound generating devices for audible presentationof output. Examples of input devices that can be used for a userinterface include keyboards, and pointing devices, such as mice, touchpads, and digitizing tablets. As another example, a computer may receiveinput information through speech recognition or in other audibleformats.

Such computers may be interconnected by one or more networks in anysuitable form, including a local area network or a wide area network,such as an enterprise network, and intelligent network (IN) or theInternet. Such networks may be based on any suitable technology and mayoperate according to any suitable protocol and may include wirelessnetworks, wired networks or fiber optic networks.

Also, as described, some aspects may be embodied as one or more methods.The acts performed as part of the method may be ordered in any suitableway. Accordingly, embodiments may be constructed in which acts areperformed in an order different than illustrated, which may includeperforming some acts simultaneously, even though shown as sequentialacts in illustrative embodiments.

All definitions, as defined and used herein, should be understood tocontrol over dictionary definitions, definitions in documentsincorporated by reference, and/or ordinary meanings of the definedterms.

The indefinite articles “a” and “an,” as used herein in thespecification and in the claims, unless clearly indicated to thecontrary, should be understood to mean “at least one.”

The phrase “and/or,” as used herein in the specification and in theclaims, should be understood to mean “either or both” of the elements soconjoined, i.e., elements that are conjunctively present in some casesand disjunctively present in other cases. Multiple elements listed with“and/or” should be construed in the same fashion, i.e., “one or more” ofthe elements so conjoined. Other elements may optionally be presentother than the elements specifically identified by the “and/or” clause,whether related or unrelated to those elements specifically identified.Thus, as a non-limiting example, a reference to “A and/or B”, when usedin conjunction with open-ended language such as “comprising” can refer,in one embodiment, to A only (optionally including elements other thanB); in another embodiment, to B only (optionally including elementsother than A); in yet another embodiment, to both A and B (optionallyincluding other elements); etc.

As used herein in the specification and in the claims, the phrase “atleast one,” in reference to a list of one or more elements, should beunderstood to mean at least one element selected from any one or more ofthe elements in the list of elements, but not necessarily including atleast one of each and every element specifically listed within the listof elements and not excluding any combinations of elements in the listof elements. This definition also allows that elements may optionally bepresent other than the elements specifically identified within the listof elements to which the phrase “at least one” refers, whether relatedor unrelated to those elements specifically identified. Thus, as anon-limiting example, “at least one of A and B” (or, equivalently, “atleast one of A or B,” or, equivalently “at least one of A and/or B”) canrefer, in one embodiment, to at least one, optionally including morethan one, A, with no B present (and optionally including elements otherthan B); in another embodiment, to at least one, optionally includingmore than one, B, with no A present (and optionally including elementsother than A); in yet another embodiment, to at least one, optionallyincluding more than one, A, and at least one, optionally including morethan one, B (and optionally including other elements); etc.

Also, the phraseology and terminology used herein is for the purpose ofdescription and should not be regarded as limiting. The use of“including,” “comprising,” or “having,” “containing,” “involving,” andvariations thereof herein, is meant to encompass the items listedthereafter and equivalents thereof as well as additional items.

In the claims, as well as in the specification above, all transitionalphrases such as “comprising,” “including,” “carrying,” “having,”“containing,” “involving,” “holding,” “composed of,” and the like are tobe understood to be open-ended, i.e., to mean including but not limitedto. Only the transitional phrases “consisting of” and “consistingessentially of” shall be closed or semi-closed transitional phrases,respectively.

What is claimed is:
 1. A low-field magnetic resonance imaging systemcomprising: a magnetics system having a plurality of magneticscomponents configured to produce magnetic fields for performing magneticresonance imaging, the magnetics system comprising: a B₀ magnetconfigured to produce a B₀ magnetic field for the low-field magneticresonance imaging system; a plurality of gradient coils configured to,when operated, generate magnetic fields to provide spatial encoding ofmagnetic resonance signals; and at least one radio frequency coilconfigured to, when operated, transmit radio frequency signals to afield of view of the low-field magnetic resonance imaging system and torespond to magnetic resonance signals emitted from the field of view;and a power system comprising: one or more power components configuredto provide power to the magnetics system to operate the low-fieldmagnetic resonance imaging system to perform image acquisition; and apower connection configured to connect to a single-phase outlet toreceive mains electricity and deliver the mains electricity to the powersystem to provide power needed to operate the low-field magneticresonance imaging system, wherein the power system is configured tooperate the low-field magnetic resonance imaging system only using powersupplied from the single-phase outlet during image acquisition.
 2. Thelow-field magnetic resonance imaging system of claim 1, wherein thepower connection is configured to connect to a single-phase outletproviding approximately between 110 and 120 volts and rated at least to30 amperes, and wherein the power system is capable of providing thepower to operate the magnetic resonance imaging system from powerprovided by the single-phase outlet.
 3. The low-field magnetic resonanceimaging system of claim 1, wherein the power connection is configured toconnect to a single-phase outlet providing approximately between 110 and120 volts and rated at least to 20 amperes, and wherein the power systemis capable of providing the power to operate the magnetic resonanceimaging system from power provided by the single-phase outlet.
 4. Thelow-field magnetic resonance imaging system of claim 1, wherein thepower connection is configured to connect to a single-phase outletproviding approximately between 110 and 120 volts and rated at least to15 amperes, and wherein the power system is capable of providing thepower to operate the magnetic resonance imaging system from powerprovided by the single-phase outlet.
 5. The low-field magnetic resonanceimaging system of claim 1, wherein the power connection is configured toconnect to a single-phase outlet providing approximately between 220 and240 volts and rated at least to 30 amperes, and wherein the power systemis capable of providing the power to operate the magnetic resonanceimaging system from power provided by the single-phase outlet.
 6. Thelow-field magnetic resonance imaging system of claim 1, wherein thepower connection is configured to connect to a single-phase outletproviding approximately between 220 and 240 volts and rated at least to20 amperes, and wherein the power system is capable of providing thepower to operate the magnetic resonance imaging system from powerprovided by the single-phase outlet.
 7. The low-field magnetic resonanceimaging system of claim 1, wherein the power connection is configured toconnect to a single-phase outlet providing approximately between 220 and240 volts and rated at least to 15 amperes, and wherein the power systemis capable of providing the power to operate the magnetic resonanceimaging system from power provided by the single-phase outlet.
 8. Thelow-field magnetic resonance imaging system of claim 1, wherein thepower connection is configured to connect to a single-phase outletproviding approximately between 220 and 240 volts and rated at least to12 amperes, and wherein the power system is capable of providing thepower to operate the magnetic resonance imaging system from powerprovided by the single-phase outlet.
 9. The low-field magnetic resonanceimaging system of claim 1, wherein the power system comprises at leastone power supply configured to receive alternating current (AC) powerfrom a wall outlet in a range between 85 volts and 250 volts atapproximately 50-60 Hertz and convert the AC power to direct current(DC) power to operate the magnetic resonance imaging system.
 10. Thelow-field magnetic resonance imaging system of claim 9, wherein thepower system comprises at least one power supply configured to receiveAC power from a single-phase wall outlet.
 11. The low-field magneticresonance imaging system of claim 1, wherein the power system operatesthe low-field magnetic resonance imaging system using an average of lessthan 3 kilowatts during image acquisition.
 12. The low-field magneticresonance imaging system of claim 1, wherein the power system operatesthe low-field magnetic resonance imaging system using an average of lessthan 2 kilowatts during image acquisition.
 13. The low-field magneticresonance imaging system of claim 1, wherein the power system operatesthe low-field magnetic resonance imaging system using an average of lessthan 1 kilowatt during image acquisition.
 14. The low-field magneticresonance imaging system of claim 1, wherein the power system operatesthe low-field magnetic resonance imaging system using an average of lessthan 750 watts during image acquisition.
 15. The low-field magneticresonance imaging system of claim 1, wherein the B₀ magnet is configuredto produce a B₀ field for the magnetic resonance imaging system at verylow-field strength of less than 0.1 T and greater than or equal toapproximately 20 mT.
 16. The low-field magnetic resonance imaging systemof claim 1, wherein the B₀ magnet is configured to produce a B₀ fieldfor the magnetic resonance imaging system at a very low-field strengthof less than 0.1 T and greater than or equal to approximately 50 mT. 17.The low-field magnetic resonance imaging system of claim 1, wherein theB₀ magnet is configured to produce a B₀ field for the magnetic resonanceimaging system at a very low-field strength of less than or equal toapproximately 50 mT and greater than or equal to approximately 20 mT.18. The low-field magnetic resonance imaging system of claim 1, whereinthe B₀ magnet is configured to produce a B₀ field for the magneticresonance imaging system at a very low-field strength of less than orequal to approximately 20 mT and greater than or equal to approximately10 mT.
 19. The low-field magnetic resonance imaging system of claim 1,wherein the one or more power components comprises a first poweramplifier to provide power to a first gradient coil, a second poweramplifier to provide power to a second gradient coil, and a third poweramplifier to provide power to a third gradient coil that together form agradient system for the low-field magnetic resonance system, and whereinthe gradient system uses an average of less than or equal toapproximately 1 kilowatt during image acquisition.
 20. The low-fieldmagnetic resonance imaging system of claim 19, wherein the gradientsystem uses an average of less than or equal to approximately 500 wattsduring image acquisition.
 21. The low-field magnetic resonance imagingsystem of claim 19, wherein the gradient system uses an average of lessthan or equal to approximately 200 watts during image acquisition. 22.The low-field magnetic resonance imaging system of claim 19, wherein thegradient system uses an average of less than or equal to approximately100 watts during image acquisition.
 23. The low-field magnetic resonanceimaging system of claim 19, further comprising a base supporting the B₀magnet, the first gradient coil, the second gradient coil and the thirdgradient coil, and housing the first power amplifier, the second poweramplifier, and the third power amplifier.
 24. The low-field magneticresonance imaging system of claim 1, wherein the one or more powercomponents comprises a first power amplifier to provide power to atleast one radio frequency coil to form a radio frequency transmit systemfor the low-field magnetic resonance system, and wherein the radiofrequency transmit system uses an average of less than or equal toapproximately 250 watts during image acquisition.
 25. The low-fieldmagnetic resonance imaging system of claim 24, wherein the radiofrequency transmit system uses an average of less than or equal toapproximately 100 watts during image acquisition.
 26. The low-fieldmagnetic resonance imaging system of claim 24, wherein the radiofrequency transmit system uses an average of less than or equal toapproximately 50 watts during image acquisition.
 27. The low-fieldmagnetic resonance imaging system of claim 23, further comprising aconveyance mechanism that allows the low-field magnetic resonanceimaging system to be transported to desired locations.
 28. A low-fieldmagnetic resonance imaging system comprising: a magnetics system havinga plurality of magnetics components configured to produce magneticfields for performing magnetic resonance imaging, the magnetics systemcomprising: a B₀ magnet configured to produce a B₀ field for thelow-field magnetic resonance imaging system; a plurality of gradientcoils configured to, when operated, generate magnetic fields to providespatial encoding of emitted magnetic resonance signals; and at least oneradio frequency coil configured to, when operated, transmit radiofrequency signals to a field of view of the low-field magnetic resonanceimaging system and to respond to magnetic resonance signals emitted fromthe field of view; and a power system comprising one or more powercomponents configured to provide power to the magnetics system tooperate the low-field magnetic resonance imaging system to perform imageacquisition only using an average number of watts less than or equal toa number of watts provided by a single-phase outlet.